Hybrid terrain-adaptive lower-extremity systems

ABSTRACT

Hybrid terrain-adaptive lower-extremity apparatus and methods that perform in a variety of different situations by detecting the terrain that is being traversed, and adapting to the detected terrain. In some embodiments, the ability to control the apparatus for each of these situations builds upon five basic capabilities: (1) determining the activity being performed; (2) dynamically controlling the characteristics of the apparatus based on the activity that is being performed; (3) dynamically driving the apparatus based on the activity that is being performed; (4) determining terrain texture irregularities (e.g., how sticky is the terrain, how slippery is the terrain, is the terrain coarse or smooth, does the terrain have any obstructions, such as rocks) and (5) a mechanical design of the apparatus that can respond to the dynamic control and dynamic drive.

RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.13/832,491, filed on Mar. 15, 2013, now U.S. Pat. No. 9,211,201, whichis a continuation of U.S. patent application Ser. No. 12/552,036 filedon Sep. 1, 2009, now U.S. Pat. No. 8,419,804, which claims priority toU.S. Provisional Patent Application Ser. No. 61/094,125, filed on Sep.4, 2008, Ser. No. 61/161,999, filed on Mar. 20, 2009 and Ser. No.61/231,754, filed on Aug. 6, 2009, the entire contents of eachapplication are hereby incorporated by reference in their entirety.

TECHNICAL FIELD

This invention relates generally to lower-extremity prosthetic, orthoticand exoskeleton apparatus, components thereof, and methods forcontrolling the same.

BACKGROUND

Over 100,000 people in the United States lose their feet throughamputation every year. Many hundreds of thousands suffer thisdebilitating loss world-wide. Additionally, there are 700,000individuals that survive a stroke each year in the United States oftencausing a variety of lower limb pathologies that constrain ambulation.Until recently, lower-extremity prosthetic and orthotic systems haveemployed predominantly passive or low-power mechanisms incapable ofdelivering the non-conservative positive work on each stride that thebody needs to achieve an economical walking motion even on flatterrain—let alone on uneven surfaCes such as stairs and steps.

It is helpful to understand the normal biomechanics associated with agait cycle of a subject to appreciate the requirements oflower-extremity prosthetic, orthotic and exoskeleton apparatus.Specifically, the function of human ankle under sagittal plane rotationis described below for different locomotor conditions.

The mechanical characteristics of conventional passive ankle/footprostheses (“AFPs”) like the Ossur Flex-Foot® remain essentiallyconstant throughout the life of the device. U.S. Patent PublishedApplication No. US 2006/0249315 (“the '315 application”) represented asignificant advance over those conventional AFPs. The '315 application,the entire contents of which are hereby incorporated by reference in itsentirety, recognized that performance can be improved by dividing thewalking cycle into five phases, and by optimizing the mechanicalcharacteristics of the device independently for each of those fivephases.

FIG. 1A. is a schematic illustration of the different phases of asubject's gait cycle over level ground. The gait cycle is typicallydefined as beginning with the heel strike of one foot and ending at thenext heel strike of the same foot. The gait cycle is broken down intotwo phases: the stance phase (about 60% of the gait cycle) and thesubsequent swing phase (about 40% of the gait cycle). The swing phaserepresents the portion of the gait cycle when the foot is off theground. The stance phase begins at heel-strike when the heel touches thefloor and ends at toe-off when the same foot rises from the groundsurface. The stance phase is separated into three sub-phases: ControlledPlantarflexion (CP), Controlled Dorsiflexion (CD), and PoweredPlantarflexion (PP).

CP begins at heel-strike illustrated at 102 and ends at foot-flat at106. CP describes the process by which the heel and forefoot initiallymake contact with the ground. Researchers have shown that that CP anklejoint behavior is consistent with a linear spring response where jointtorque is proportional to the displacement of the joint in relation toan equilibrium position of the joint position. The spring behavior is,however, variable; joint stiffness is continuously modulated by the bodyfrom step to step within the three sub-phases of stance and late swingstate.

After the CP period, the CD phase continues until the ankle reaches astate of maximum dorsiflexion and begins powered plantarflexion PP asillustrated at 110. Ankle torque versus position during the CD period isdescribed as a nonlinear spring where stiffness increases withincreasing ankle position. The ankle stores the elastic energy during CDwhich is necessary to propel the body upwards and forwards during the PPphase.

The PP phase begins after CD and ends at the instant of toe-offillustrated at 114. During PP, the ankle applies torque in accordancewith a reflex response that catapults the body upward and forward. Thecatapult energy is then released along with the spring energy storedduring the CD phase to achieve the high plantarflexion power during latestance. This catapult behavior is necessary because the work generatedduring PP is more than the negative work absorbed during the CP and CDphases for moderate to fast walking speeds. The foot is lifted off theground during the swing phase, from toe-off at 114 until the next heelstrike at 118.

Because the kinematic and kinetic patterns at the ankle during stairascent/descent are different from that of level-ground walking, aseparate description of the ankle-foot biomechanics is presented inFIGS. 1B and 1C. FIG. 1B shows the human ankle biomechanics during stairascent. The first phase of stair ascent is called ControlledDorsiflexion 1 (CD 1), which begins with foot strike in a dorsiflexedposition seen at 130 and continues to dorsiflex until the heel contactsthe step surface at 132. In this phase, the ankle can be modeled as alinear spring. The second phase is Powered Plantar flexion 1 (PP 1),which begins at the instant of foot flat (when the ankle reaches itsmaximum dorsiflexion at 132) and ends when dorsiflexion begins onceagain at 134. The human ankle behaves as a torque actuator to provideextra energy to support the body weight.

The third phase is Controlled Dorsiflexion 2 (CD 2), in which the ankledorsiflexes until heel-off at 136. For the CD 2 phase, the ankle can bemodeled as a linear spring. The fourth and final phase is PoweredPlantar flexion 2 (PP 2) which begins at heel-off 136 and continues asthe foot pushes off the step, acting as a torque actuator in parallelwith the CD 2 spring to propel the body upwards and forwards, and endswhen the toe leaves the surface at 138 to begin the swing phase thatends at 140.

FIG. 1C shows the human ankle-foot biomechanics for stair descent. Thestance phase of stair descent is divided into three sub-phases:Controlled Dorsiflexion 1 (CD 1), Controlled Dorsiflexion 2 (CD2), andPowered Plantar flexion (PP). CD1 begins at foot strike illustrated at150 and ends at foot-flat 152. In this phase, the human ankle can bemodeled as a variable damper. In CD2, the ankle continues to dorsiflexforward until it reaches a maximum dorsiflexion posture seen at 154.Here the ankle acts as a linear spring, storing energy throughout CD2.During PP, which begins at 154, the ankle plantar flexes until the footlifts from the step at 156. In this final PP phase, the ankle releasesstored CD2 energy, propelling the body upwards and forwards. Aftertoe-off at 156, the foot is positioned controlled through the swingphase until the next foot strike at 158.

For stair ascent depicted in FIG. 1B, the human ankle-foot can beeffectively modeled using a combination of an actuator and a variablestiffness mechanism. However, for stair descent, depicted in FIG. 1C, avariable damper needs also to be included for modeling the ankle-footcomplex; the power absorbed by the human ankle is much greater duringstair descent than the power released during stair ascent. Hence, it isreasonable to model the ankle as a combination of a variable-damper andspring for stair descent.

Conventional passive prosthetic, orthotic and exoskeleton apparatus donot adequately reproduce the biomechanics of a gait cycle. They are notbiomimetic because they do not actively modulate impedance and do notapply the reflexive torque response; neither on level ground, ascendingor descending stairs or ramps, or changing terrain conditions. A needtherefore exists for improved lower-extremity prosthetic, orthotic andexoskeleton apparatus, components thereof, and methods for controllingthe same.

SUMMARY

The inventors have recognized that during the course of an ordinary day,a person's lower limbs are used to perform and adapt to many differentactivities in addition to ordinary walking, such as ascending anddescending stairs, and walking on inclined ramps. The ankle-footcomponents require the most power and must exhibit the mostterrain-adaptive behavior because these are in the most direct contactwith the underlying terrain. The inventors have further recognized thatthe performance of AFPs can be dramatically improved by dynamicallyoptimizing the mechanical characteristics of the device in differentways and dynamically controlling the device in different ways for eachof those activities.

For example, when a person is walking on flat ground, it is better tocontrol the angle of the foot so that the heel is lower than the toewhen the foot touches down on the ground. However, when a person isascending stairs, it is better to control the angle of the foot so thatthe toe is lower than the heel when the foot touches down on the nextstep.

This application describes various embodiments of AFPs that performappropriately in each of these different situations by detecting theterrain that is being traversed, and automatically adapting to thedetected terrain. In some embodiments, the ability to control the AFPfor each of these situations builds upon five basic capabilities: (1)determining the activity being performed; (2) dynamically controllingthe characteristics of the AFP based on the activity that is beingperformed; (3) dynamically driving the AFP based on the activity that isbeing performed; (4) determining terrain texture irregularities (e.g.,how sticky is the terrain, how slippery is the terrain, is the terraincoarse or smooth, does the terrain have any obstructions, such as rocks)and responding to these with appropriate fraction control and (5) amechanical design of the AFP that can respond to the dynamic control anddynamic drive.

The inventors have determined that an exemplary way to figure out whatactivity is being performed is to track the trajectory of a spot(typically at the virtual center of rotation of the ankle joint) on thelower leg (or shank) between the ankle joint and knee joint. FIG. 6Ashows the shank trajectories that correspond to five differentactivities, with additional ramp trajectories to distinguish betweensteep and shallow ramps. The system can use this information to figureout what activity is being performed by mapping the tracked trajectoryonto a set of activities.

By looking at the trajectory of the lower leg (shank) it is possible todistinguish between flat terrain, ascending or descending stairs, orascending or descending ramps. For example, when the system recognizes atrajectory it would switch into an appropriate mode, and dynamicallycontrol (drive) the AFP as previously established for the mode. Where atrajectory does not fall neatly within a classification, the AFPcontroller would optimize the response to minimize an objective functionin a stochastic control sense or would apply fuzzy logic or adhoccontrols based upon the likelihood the terrain falls into aclassification.

One suitable way to track the trajectory of the shank is by mounting aninertial measurement unit (IMU) at the forward face at the top of thelower leg member (shank), and processing the signals that are output bythe IMU. A suitable way to distinguish the various trajectories is tomonitor the velocity of the ankle joint angle of attack. These topicsare described in greater detail below.

In addition to dynamically optimizing the mechanical characteristics anddynamically controlling the device in different ways for each of thedifferent activities, the inventors have recognized that the performanceof the device can be further improved by fine-tuning the characteristicsand control of the AFP based on various parameters.

For example, when a person is walking slowly (e.g., at a rate of lessthan 0.9 meters per second), performance can be improved by increasingthe impedance of the ankle joint with respect to the impedance used fornormal walking. Or when a person is walking quickly (e.g., at a rate of1.7 meters per second), performance can be improved by decreasing theimpedance of the ankle joint with to the impedance used for normalwalking.

In addition, when the controller determines that the ankle joint is notresponding as we would expect it to when traversing normal terrain, thecontroller can take into account (and modify the output of thecontroller) that there may be features, texture or irregularities in theterrain (e.g., how sticky is the terrain, how slippery is the terrain,is the terrain coarse or smooth, does the terrain have any obstructions,such as rocks).

Each of the five capabilities identified above (i.e., figuring out whatactivity is being performed; figuring out whether there are features,texture or irregularities of the terrain; dynamically controlling thecharacteristics of the AFP; dynamically driving the AFP; and themechanical design of the AFP) is described in greater detail below.

The inventions described herein relate generally to lower-extremityprosthetic, orthotic and exoskeleton apparatus. Typical use cases forvarious embodiments of the invention include, for example, metabolicaugmentation, permanent assistance for subjects with a permanent limbpathology, or rehabilitation for wearers with temporary limb pathology.

An example of a use case for an exemplary lower-extremity prostheticapparatus (e.g., apparatus 1700 of FIGS. 17A-17G) involves theprosthetic replacing the ambulation function of a lower limb of thewearer. An example of a use case for an exemplary lower-extremityorthotic apparatus (e.g., apparatus 2200 of FIGS. 22A-22C—orthosis)features a method for determining a level of assistance desired forapparatus to apply to a wearer wearing the apparatus. In someembodiments, the level of assistance performed by the orthosis isreduced based on impedance and torque contribution of the wearer to theapparatus. In some embodiments, the impedance and torque contribution ofthe wearer is determined based on a dynamic, biomechanical model of thewearer and apparatus and measurements of the wearer during operation ofthe apparatus. In some embodiments, the measurements of the wearerinclude at least one of rotation and acceleration of at least one jointof the apparatus. In some embodiments, the axial force and momentapplied to the lower leg member of the apparatus is determined based onsensor measurements made using a structural member (pyramid) coupled tothe lower leg member of the apparatus. The pyramid is an instrumentedstructure that is a component of a prosthesis and which couples to thelimb socket of the wearer. In one embodiment, the pyramid (structuralelement) measurements are used by a controller to determine axial forceand moment applied to the lower leg member. In some embodiments, theapparatus includes at least one of an ankle joint that connects a footmember of the apparatus to a lower leg member of the apparatus, or aknee joint for connecting a thigh member of the apparatus to the lowerleg member of the apparatus, or a hip joint for connecting a torsomember of the apparatus to the thigh member of the apparatus.

The invention, in one aspect, features a low noise linear actuator thatincludes a rotary motor comprising a motor shaft output. The actuatoralso includes a screw transmission assembly that includes a threadedshaft coupled to the motor shaft output, the threaded shaft comprising ahollowed out portion containing an acoustic damping material and a nutassembly. The screw transmission assembly translates rotational motionof the motor shaft output to a linear motion of the nut assembly.

In some embodiments, the screw transmission assembly is a ball-screwtransmission assembly and the nut assembly is a ball-nut assembly,wherein the ball-nut assembly also includes a plurality of ball bearingsand a plurality of ball tracks for holding the ball bearings and forrecirculating the ball bearings in the ball bearing assembly. In someembodiments, the actuator includes a pulley coupling the motor shaftoutput to the threaded shaft via a plurality of belts connected inparallel between the pulley and the threaded shaft of the ball-screwtransmission assembly. In some embodiments, the linear actuator includesa sensor that validates belt integrity during operation. The pulley canbe welded to the motor shaft output.

In some embodiments, the actuator includes a radial and thrust bearingcoupling the plurality of belts to the threaded shaft to support loadsapplied by tension in the belts and the threaded shaft. In someembodiments, the ball-screw transmission assembly includes at least oneseal for protecting the ball-screw transmission assembly fromcontaminants. In some embodiments, the linear actuator is a component ofa lower extremity prosthesis orthosis, or exoskeleton. In someembodiments, the linear actuator includes a transmission that employstraction wheels that couple the motor shaft output to the threaded shaftof the ball-screw transmission assembly. The screw transmission assemblycan be a lead screw transmission assembly.

The invention, in another aspect, features a linear actuator thatincludes a rotary motor comprising a motor shaft output and a motordrive transmission assembly coupled to the motor shaft output totranslate rotational motion of the motor shaft output to a linear motionat an output of the motor drive transmission. The linear actuator alsoincludes at least one elastic element with bi-directional stiffnessconnected in series with the motor drive transmission assembly to storeenergy in tension and compression.

In some embodiments, the linear actuator includes a strain sensorcoupled to the at least one elastic element for measuring strains in theat least one elastic element. The at least one elastic element can be aseries or parallel elastic element coupled to the output of the motordrive transmission assembly. In some embodiments, the linear actuatorincludes a controller for receiving measured strain signals forperforming closed loop control of the linear actuator thrust force. Theat least one elastic element can be a substantially flat spring dividedalong a longitudinal axis of the spring minimizing out-of-plane momentapplied by the spring to the output of the motor drive transmissionassembly. The at least one elastic element can be a series elasticstrain element coupled to the output of the motor drive transmissionassembly, and the linear actuator can also include a sensor thatmeasures motor position or position of the output of the motor drivetransmission assembly, and at least one sensor that measures the outputof the series elastic element, and signal processing electronics thatestimates thrust force of the linear actuator for closed loop control ofthe linear actuator thrust force.

The invention, in another aspect, features a lower-extremity prosthesis,orthosis or exoskeleton apparatus that includes a foot member, a lowerleg member and an ankle joint for connecting the foot member to thelower leg member. The apparatus also includes a first actuator forapplying torque to the ankle joint to rotate the foot member withrespect to the lower leg member. The apparatus also includes at leastone passive elastic members that is a non-compliant stop connected inparallel with the actuator between the lower leg member and the footmember, wherein the non-compliant stop stores little or no energy duringdorsiflexion and limits further rotation of the ankle beyond apredefined angle during powered plantar flexion.

In some embodiments, the apparatus includes an angle adjustmentmechanism for setting a pre-specified angle of the foot member relativeto the lower leg member at which the non-compliant stop limits furtherrotation. The angle adjustment mechanism can include a screw adjustablecomponent for setting the pre-specified angle. The angle adjustmentmechanism can include an actuator for setting the pre-specified angle.In some embodiments, the actuator adjusts the pre-specified angle basedon a property of the underlying terrain. In some embodiments, theproperty of the underlying terrain is selected from the group consistingof ascending ramp, descending ramp, ascending stair, descending stair,level surface. In some embodiments, a controller associated with theapparatus determines the property of the underlying terrain on anintra-cycle basis.

The invention, in another aspect, features a lower-extremity prosthesisthat includes a foot member, a lower leg member, and an ankle joint forconnecting the foot member to the lower leg member. The prosthesisincludes a first actuator for applying torque to the ankle joint torotate the foot member with respect to the lower leg member. Theprosthesis also includes a structural element coupled to the lower legmember and comprising an interface for coupling to a limb socket memberof a wearer, wherein the structural element comprises a plurality ofstrain gages for determining the torque applied to the lower leg memberby the actuator and the axial force applied to the lower leg member.

In some embodiments, the prosthesis includes an inertial measurementunit for determining an inertial pose of the lower leg member. Theinertial measurement unit can be coupled to the lower leg member. Theinertial measurement unit can be coupled to the foot member. In someembodiments, the prosthesis includes a controller for calculating groundreaction force and zero moment pivot coordinates imparted by anunderlying surface onto the foot member based on an inertial pose of thelower leg member, the torque applied to the lower leg member by theactuator, axial force applied to the lower leg member, and an anglebetween the foot member and lower leg member. In some embodiments, thecontroller is coupled to the actuator and is configured to control theactuator for modulating at least one of an impedance, position or torqueof the prosthesis throughout a walking cycle of the prosthesis based onthe inertial pose trajectory of the lower leg member, the angle betweenthe foot member and lower leg member, and the ground reaction force andthe zero moment pivot coordinates. In some embodiments, the controlleris coupled to the actuator and is configured to control the actuator formodulating an impedance of the prosthesis as the wearer stands up from aseated position or sits down from a standing position based on theinertial pose of the lower leg member, the angle between the foot memberand lower leg member, and the ground reaction force and the zero momentpivot coordinates.

The invention, in another aspect, features a lower-extremity prosthesisthat includes a foot member, a lower leg member and an ankle joint forconnecting the foot member to the lower leg member. The prosthesisincludes a first actuator for applying torque to the ankle joint torotate the foot member with respect to the lower leg member. Theprosthesis also includes a structural element coupled to the lower legmember and includes an interface for coupling to a limb socket member ofa wearer. The prosthesis also includes a displacement sensing apparatusfor measuring deflection of the structural element for determining thetorque applied to the lower leg member by the actuator and the axialforce applied to the lower leg member.

In some embodiments, the displacement sensing apparatus includes aplurality of sensors and the displacement sensing apparatus measures thedistance between each sensor and a surface of the structural element.The sensors can be selected from the group consisting of contactdisplacement sensors, non-contact displacement sensors, inductive coils,optical sensors, force-sensitive resistors, piezoelectric sensors, orstrain sensors. In some embodiments, the plurality of sensors include aplurality of inductive coils on a circuit board. In some embodiments,changes in inductance of the inductive coils relative to a surface ofthe structural element are used to determine the displacement of thestructural element.

In some embodiments, the prosthesis includes an inertial measurementunit for determining an inertial pose of the lower leg member. In someembodiments, the inertial measurement unit is coupled to the lower legmember. In some embodiments, the inertial measurement unit is coupled tothe foot member.

In some embodiments, the prosthesis includes a controller forcalculating ground reaction force and zero moment pivot coordinatesimparted by an underlying surface onto the foot member based on aninertial pose trajectory of the lower leg member, the torque applied tothe lower leg member by the actuator, axial force applied to the lowerleg member, and an angle between the foot member and lower leg member.In some embodiments, the controller is coupled to the actuator and isconfigured to control the actuator for modulating an impedance of theprosthesis throughout a walking cycle of the prosthesis based on theinertial pose trajectory of the lower leg member, the angle between thefoot member and lower leg member, and the ground reaction force and thezero moment pivot coordinates. In some embodiments, the controller iscoupled to the actuator and is configured to control the actuator formodulating an impedance of the prosthesis as the wearer stands up from aseated position or sits down from a standing position based on theinertial pose trajectory of the lower leg member, the angle between thefoot member and lower leg member, and the ground reaction force.

The invention, in another aspect, features an active knee orthosis thatincludes a thigh member attachable to a thigh of a wearer, a lower legmember attachable to a lower leg of the wearer and a knee joint forconnecting the thigh member to the lower leg member. The orthosis alsoincludes a rotary motor comprising a motor shaft output. The orthosisalso includes a motor drive transmission assembly coupled to the motorshaft output to translate rotational motion of the motor shaft output toa linear motion at an output of the motor drive transmission assembly.The orthosis also includes a drive transmission assembly coupled to theoutput of the motor drive transmission, an output of the drivetransmission assembly is coupled to the lower leg member for applyingtorque to the knee joint to rotate the lower leg member with respect tothe thigh member. The orthosis also includes a motor angle sensor fordetermining motor position. The orthosis also includes a controller forcontrolling the rotary motor for modulating impedance, position ortorque of the of the orthosis throughout a walking cycle of the orthosisbased on the motor position.

In some embodiments, the orthosis includes an angle sensor fordetermining position of a drum of the drive transmission assemblyrelative to the output of the motor drive transmission assembly andwherein the controller controls the rotary motor for modulatingimpedance, position or torque based on the position. In someembodiments, the orthosis includes a displacement sensor for measuringdisplacement of a series spring in the motor drive transmission assemblyfor determining force on the series spring and wherein the controllercontrols the rotary motor for modulating impedance, position or torquebased on the force on the spring.

In some embodiments, the orthosis includes an inertial measurement unitcoupled to the thigh member or lower leg member for determining aninertial pose trajectory of the lower leg member and wherein thecontroller controls the rotary motor for modulating impedance, positionor torque based on the inertial pose. In some embodiments, the orthosisincludes a sensor for determining the force applied to at least one ofthe lower leg member and thigh member by the drive transmission assemblyand wherein the controller controls the rotary motor for modulatingimpedance, position or torque based on the torque applied to the lowerleg member.

In some embodiments, the orthosis includes an angle sensor fordetermining an angle between the thigh member and lower leg member andwherein the controller controls the rotary motor for modulatingimpedance, position or torque based on the angle between the thighmember and lower leg member. In some embodiments, the orthosis includesthe drive transmission is selected from the group consisting of a beltdrive transmission, band drive transmission and cable drivetransmission. In some embodiments, the orthosis includes a cuff coupledto the thigh member for attaching the orthosis to the thigh of thewearer. In some embodiments, the orthosis includes a cuff coupled to thelower leg member for attaching the orthosis to the lower leg of thewearer. In some embodiments, the orthosis augments lower extremityfunctions of the wearer. In some embodiments, the orthosis treats alower extremity pathology of the wearer. In some embodiments, thecontroller is configured to vary assistance provided by the orthosis tothe wearer during rehabilitation of a lower extremity pathology of thewearer.

The invention, in another aspect, features an active knee orthosis thatincludes a thigh member attachable to a thigh of a wearer, a lower legmember attachable to a lower leg of the wearer, and a knee joint forconnecting the thigh member to the lower leg member. The orthosis alsoincludes a rotary motor comprising a motor shaft output. The orthosisalso includes a screw transmission assembly coupled to the motor shaftoutput for converting the rotary motion of the motor shaft output to alinear motion output by the screw transmission assembly. The orthosisalso includes a belt, band or cable drive transmission assembly coupledto the output of the screw transmission assembly to convert a linearmotion output by the screw transmission assembly to a rotary motion forapplying torque to the knee joint to rotate the lower leg member withrespect to the thigh member. The orthosis also includes a motor anglesensor for determining motor position. The orthosis also includes acontroller for controlling the rotary motor for modulating impedance,position or torque of the of the orthosis throughout a walking cycle ofthe orthosis based on the motor position.

In some embodiments, the orthosis includes a displacement sensor formeasuring displacement of a series spring in the belt, band or cabledrive transmission for determining force on the series spring andwherein the controller controls the rotary motor for modulatingimpedance, position or torque based on the force on the spring. In someembodiments, the orthosis includes an inertial measurement unit coupledto the thigh member or lower leg member for determining, within a gaitcycle, an inertial pose trajectory of the lower leg member and whereinthe controller controls the rotary motor for modulating impedance,position or torque based on the inertial pose trajectory during the gaitcycle.

In some embodiments, the orthosis includes a torque sensor fordetermining torque applied to the lower leg member by the belt, band orcable drive transmission and wherein the controller controls the rotarymotor for modulating impedance, position or torque within the gait cyclebased on the force applied to the lower leg member. In some embodiments,the orthosis includes an angle sensor for determining an angle betweenthe thigh member and lower leg member and wherein the controllercontrols the rotary motor for modulating impedance, position or torquebased on the angle between the thigh member and lower leg member withinthe gait cycle. In some embodiments, the belt, band or cable drivetransmission comprises at least two drive transmissions, wherein a firstof the at least two drive transmissions is configured to convert a firstdirection of a linear motion output by the screw transmission assemblyto a first rotary motion for applying torque to the knee joint to rotatethe lower leg member with respect to the thigh member and wherein asecond of the at least two transmissions is configured to convert anopposite direction of a linear motion output by the screw transmissionassembly to an opposite rotary motion for applying torque to the kneejoint to rotate the lower leg member with respect to the thigh member.

The invention, in another aspect, features a method for determiningground reaction forces and zero moment pivot imparted by an underlyingsurface onto a foot member of a lower extremity prosthetic apparatusworn by a wearer. The apparatus includes a foot member, a lower legmember, and an ankle joint for connecting the foot member to the lowerleg member and a first actuator for applying torque to the ankle jointto rotate the foot member with respect to the lower leg member. Themethod involves calculating the ground reaction force based on aninertial pose of the lower leg member, the torque applied to the lowerleg member by the actuator, axial force applied to the lower leg member,and an angle between the foot member and lower leg member.

In some embodiments, the method includes controlling the actuator formodulating an impedance of the apparatus throughout a walking cycle ofthe apparatus based on the inertial pose of the lower leg member, theangle between the foot member and lower leg member, the ground reactionforce and the zero moment pivot. In some embodiments, the methodincludes controlling the actuator for modulating an impedance of theapparatus as the wearer stands up from a seated position or sits downfrom a standing position based on the inertial pose of the lower legmember, the angle between the foot member and lower leg member, theground reaction force and the zero moment pivot. In some embodiments,the inertial pose of the lower leg member is determined based on anoutput of an inertial measurement unit coupled to the lower leg member.

The invention, in another aspect, features a method for minimizing theeffect of accelerometer and rate gyro errors on a lower extremityprosthesis or orthosis that includes a foot member, a lower leg member,and an ankle joint for connecting the foot member to the lower legmember. The method includes determining at least one velocity errorcontribution for an accelerometer signal output by an accelerometercoupled to the lower leg member when the ankle joint is substantiallystationary during a walking cycle of the prosthesis or orthosis. Themethod also includes determining at least one velocity errorcontribution for an inertial pose misalignment signal output by aninertial measurement unit coupled to the lower leg member when the anklejoint is substantially stationary during a walking cycle of theprosthesis or orthosis.

In some embodiments, the inertial pose misalignment signal output by theinertial measurement unit is a rate gyro signal output by a rate gyro.In some embodiments, the method includes computing the pose of the lowerleg member using signals output by the accelerometer and rate gyro. Insome embodiments, the method includes correcting the computed pose ofthe lower leg member using the velocity error contributions. In someembodiments, the method includes the velocity error contributions aredetermined during a portion of a controlled dorsiflexion state of thewalking cycle.

In some embodiments, the method includes determining velocity errorcontributions for an accelerometer signal and rate gyro signal output byan accelerometer and rate gyro coupled to a thigh member of theprosthesis or orthosis when the ankle joint is substantially stationaryduring a walking cycle of the prosthesis or orthosis. In someembodiments, the method includes determining velocity errorcontributions for an accelerometer signal and rate gyro signal output byan accelerometer and rate gyro coupled to a thigh member of theprosthesis or orthosis when a computed position on a foot member issubstantially stationary.

In some embodiments, the method includes measuring the angle of thelower leg member relative to the thigh member. In some embodiments, themethod includes determining velocity error contributions for anaccelerometer signal and rate gyro signal output by an accelerometer andrate gyro coupled to a wearer's torso when the ankle joint issubstantially stationary during a walking cycle of the prosthesis ororthosis. In some embodiments, the method includes measuring the angleof the thigh member relative to the wearer's torso.

The invention, in another aspect, features a method for controllingbalance of a wearer wearing a lower extremity prosthetic, orthotic orexoskeleton apparatus that includes a foot member, a lower leg member,and an ankle joint for connecting the foot member to the lower legmember. The method includes adjusting at least one of the ankle jointimpedance, position or torque of the apparatus based on inertial pose ofthe lower leg member, angle between the lower leg member and the footmember and ground reaction force and the zero moment pivot imparted byan underlying surface onto the foot member.

In some embodiments, the actuator coupled to the lower leg member andfoot member, adjusts the at least one of the ankle joint impedance,position or torque to control the balance of the wearer. In someembodiments, a controller calculates the ground reaction force and thezero moment pivot based on an inertial pose of the lower leg member, thetorque applied to the lower leg member by the actuator, axial forceapplied to the lower leg member, and an angle between the foot memberand lower leg member, the controller is coupled to the actuator tocontrol the actuator to adjust the at least one of the ankle jointimpedance, position or torque to control the balance of the wearer. Insome embodiments, the controller calculates the inertial pose of thelower leg based on a signal output from an inertial measurement unitcoupled to the lower leg. In some embodiments, a controller coupled tothe actuator controls the actuator to adjust the at least one of theankle joint impedance, position or torque to control the balance of thewearer. In some embodiments, the controller receives signals from one ormore sensors to calculate the inertial pose of the lower leg member,angle between the lower leg member and the foot member and the groundreaction force imparted by the underlying surface onto the foot member.

In some embodiments, the method includes controlling balance of thewearer as the wearer transitions from a sitting position to a standingposition based on an increase in the ground reaction force. In someembodiments, the method includes driving the lower leg member forwardwith an actuator coupled to the lower leg based on the increase in theground reaction force.

The invention, in another aspect, features a method for determining achange in traction between a foot member of an orthotic, prosthetic orexoskeleton apparatus and an underlying surface, the apparatus includesg a foot member, a lower leg member, an ankle joint for connecting thefoot member to the lower leg member and a first actuator for applyingtorque to the ankle joint to rotate the foot member with respect to thelower leg member. The method includes calculating ground reaction forceand the zero moment pivot imparted by an underlying surface onto thefoot member based on an inertial pose of the lower leg member, thetorque applied to the lower leg member by the actuator, axial forceapplied to the lower leg member, and an angle between the foot memberand lower leg member. The method also includes calculating velocity ofthe foot member zero moment pivot based on the inertial pose of thelower leg member, the torque applied to the lower leg member by theactuator, the axial force applied to the lower leg member, and the anglebetween the foot member and lower leg member.

In some embodiments, wherein it is determined that the foot member isslipping or sinking if the velocity of the foot member zero moment pivotdecreases during a portion of a gait cycle of the wearer between afoot-flat and toe-off condition. In some embodiments, the methodincludes reducing torque applied to the lower leg member in response todetermining that the foot member is slipping or sinking. In someembodiments, the method includes reducing the torque applied to thelower leg member by an attenuation factor. The attenuation factor can bea predetermined attenuation factor. The attenuation factor can bedetermined based on the zero moment pivot velocity. In some embodiments,the method includes reducing the torque applied to the lower leg memberin response to the zero moment pivot velocity being below apredetermined threshold.

The invention, in another aspect, features a linear actuator havingintrinsic safety features. The actuator includes a rotary motor thatincludes a motor shaft output, wherein a pulley is coupled to the motorshaft output. The actuator also includes a ball-screw transmissionassembly that includes a threaded shaft coupled to the motor shaftoutput by a plurality of belts connected in parallel between the pulleyand the threaded shaft of the ball-screw transmission assembly. Theball-screw transmission assembly translates rotational motion of themotor shaft output to a linear motion of a portion of the ball-screwtransmission assembly.

In some embodiments, the linear actuator includes an angular encoder fordetermining angular alignment between the rotary motor's rotor andstator. In some embodiments, the linear actuator includes a controllerconfigured to short three leads of the rotary motor to ground inresponse to a belt breakage sensor detecting a failure of one or more ofthe plurality of belts. In some embodiments, shorting the three leadsresults in the rotary motor functioning as a stiff, viscous brake. Insome embodiments, the temperature of the motor is determined by applyinga fixed current to a winding of the motor winding and measuring acorresponding voltage in the winding to determine the windingresistance. In some embodiments, the temperature of the motor isdetermined by applying a fixed voltage to a winding of the motor windingand measuring a corresponding current in the winding to determine thewinding resistance. In some embodiments, the linear actuator includes amotor temperature sensor for measuring the temperature of the motor. Insome embodiments, the linear actuator includes a controller coupled tothe motor for controlling torque output by the actuator based ontemperature of the motor.

The invention, in another aspect, features a method for controllingthroughout a gait cycle at least one of joint position, impedance ortorque of a lower-extremity prosthetic, orthotic, or exoskeletonapparatus worn by a wearer based on an inertially-referenced,intra-cycle trajectory of a portion of the apparatus over underlyingterrain.

In some embodiments, the apparatus includes a foot member, a lower legmember, and an ankle joint for connecting the foot member to the lowerleg member. In some embodiments, the apparatus includes a lower legmember, a thigh member, and a knee joint for connecting the lower legmember to the thigh member. In some embodiments, the apparatus includesa thigh member, a torso member, and a hip joint for connecting the thighmember to the torso member. In some embodiments, the apparatus includesa thigh member and a knee joint for connecting the lower leg member tothe thigh member. In some embodiments, the apparatus includes a torsomember, and a hip joint for connecting the thigh member to the torsomember. In some embodiments, the apparatus includes a torso member, anda hip joint for connecting the thigh member to the torso member.

In some embodiments, the trajectory is determined for the lower legmember. In some embodiments, the trajectory is determined based on aninertial pose of the lower leg member and an angle between the footmember and lower leg member. In some embodiments, the spring equilibriumposition of the foot member is adjusted to a foot-flat position relativeto the underlying terrain to coincide with the lower leg member being ina vertical position relative to a world coordinate system. In someembodiments, the impedance of the apparatus is adjusted to minimize acost function based on projected force imparted on the lower leg memberduring a period of time between when a foot member strikes theunderlying terrain and when the foot member is positioned in a flat-footposition relative to the underlying terrain. In some embodiments, theimpedance of the apparatus is adjusted to minimize a cost function basedon projected force imparted on the lower leg member during a period oftime between when a foot member strikes the underlying terrain to whenthe foot member is positioned in a flat-foot position relative to theunderlying terrain.

In some embodiments, the impedance of the apparatus is adjusted tominimize foot slap of the foot member. In some embodiments, the positionof the foot member is adjusted to a toe down position relative to theunderlying terrain based on the trajectory of the lower leg member. Insome embodiments, the trajectory of the lower leg member isrepresentative of trajectory when the underlying surface comprises oneor more stairs. In some embodiments, the at least one of joint position,impedance or torque is updated continuously during the gait cycle by aprocessor in communication with at least one sensor and one actuator ofthe apparatus. In some embodiments, the impedance and torque on thejoint of the apparatus is controlled during a late stance phase of thegait cycle based on at least one of ambulation speed, terrain context orterrain texture. In some embodiments, the impedance and torque arecontrolled to achieve a desired amount of work.

In some embodiments, the impedance of the apparatus is adjusted during acontrolled plantar flexion phase of the gait cycle to minimize forefootcollisions with the underlying terrain. In some embodiments, the atleast one of joint position, impedance or torque of the apparatus iscontrolled based on speed of a portion the apparatus.

In some embodiments, the apparatus is a lower-leg apparatus and theportion is a location between a knee joint and ankle joint of thelower-leg apparatus. In some embodiments, throughout the gait cycle, atleast two of the joint position, the impedance or the torque arecontrolled. In some embodiments, the apparatus includes, throughout thegait cycle, the joint position, the impedance and the torque arecontrolled.

The invention, in another aspect, features a method for reducing,throughout a gait cycle, hip impact force and hip impact force rate of alower-extremity prosthetic, orthotic, or exoskeleton apparatus worn by awearer. The method includes generating a cost function based on hipimpact force and force rate generated by transmission of foot contactwith underlying terrain. The method also includes controlling at leastone of position, impedance or torque of at least one joint of thelower-extremity prosthetic, orthotic, or exoskeleton apparatus based onminimizing the cost function wearer to reducing hip impact forcesgenerated during a gait cycle over the underlying terrain.

In some embodiments, the apparatus includes a first foot member, firstlower leg member and a first ankle joint for connecting the first footmember to the first lower leg member, and the method also includesadjusting impedance of the first ankle joint and an angle between thefirst foot member and first lower leg member during a time intervalbetween a foot-strike condition and foot-flat condition of the firstfoot member of the apparatus. In some embodiments, the impedance of thefirst ankle joint and the angle between the first foot member and thefirst lower leg member is adjusted to minimize a cost function based onan estimation of force to be imparted on the first ankle joint betweenthe foot-strike condition and the foot-flat condition of the first footmember of the apparatus. In some embodiments, the foot-strike conditionincludes the foot member heel first striking the underlying terrain. Insome embodiments, the foot-strike condition comprises the foot membertoe first striking the underlying terrain.

In some embodiments, the underlying terrain includes at least oneascending or descending stair, and the method also includes constrainingthe first foot member to achieve a toe first striking of the underlyingterrain while minimizing the cost function based on the estimation offorce to be imparted on the first ankle joint between the foot-strikecondition and the foot-flat condition of the first foot member of theapparatus.

In some embodiments, the apparatus includes a second leg member, asecond foot member and a second ankle joint for connecting the secondleg member to the second foot member, and the method also includesapplying a torque to the second ankle joint at or before time of impactof the first foot member with the underlying terrain. In someembodiments, the method includes controlling at least two of the jointposition, the impedance or the torque. In some embodiments, the methodincludes controlling the joint position, the impedance and the torque.

The invention, in another aspect, features a method for minimizing,throughout a gait cycle, work performed by a lower-extremity prosthetic,orthotic, or exoskeleton apparatus worn by a wearer. The method includesgenerating a cost function for estimating intra-step transition workperformed by a combination of the apparatus and subject on center ofmass of the combination during double support. The method also includescontrolling at least one of position, impedance or torque of at leastone joint of the lower-extremity prosthetic, orthotic, or exoskeletonapparatus based on minimizing the cost function wearer to reducing thework performed by the wearer and apparatus generated during a gaitcycle.

In some embodiments, the apparatus includes a first foot member, firstlower leg member and a first ankle joint for connecting the first footmember to the first lower leg member, and the method also includesadjusting impedance of the first ankle joint and an angle between thefirst foot member and first lower leg member during a time intervalbetween a foot-strike condition and foot-flat condition of the firstfoot member of the apparatus. In some embodiments, the impedance of thefirst ankle joint and the angle between the first foot member and thefirst lower leg member is adjusted to minimize a cost function based onan estimation of force to be imparted on the first ankle joint betweenthe foot-strike condition and the foot-flat condition of the first footmember of the apparatus. The foot-strike condition can include the footmember heel first striking underlying terrain. The foot-strike conditioncan include the foot member toe first striking underlying terrain.

In some embodiments, the terrain underlying the wearer includes at leastone ascending or descending stair, and the method also includesconstraining the first foot member to achieve a toe first striking ofthe underlying terrain while minimize the cost function based on theestimation of force to be imparted on the first ankle joint between thefoot-strike condition and the foot-flat condition of the first footmember of the apparatus.

In some embodiments, the apparatus includes a second leg member, asecond foot member and a second ankle joint for connecting the secondleg member to the second foot member, and the method also includesapplying a torque to the second ankle joint at or before time of impactof the first foot member with underlying terrain. In some embodiments,the method includes controlling throughout the gait cycle at least twoof the joint position, the impedance or the torque. In some embodiments,the method includes controlling throughout the gait cycle the jointposition, the impedance and the torque.

The invention, in another aspect, features a method for controlling atleast one of joint impedance, position or torque of a lower-extremityprosthetic, orthotic or exoskeleton apparatus worn by a wearer duringintra-cycle ambulation. The method also includes determining trajectoryof a location between an ankle joint and knee joint of the apparatus ina coordinate system throughout a walking cycle. The method also includesadjusting the articulation of a foot member of the apparatus based onthe trajectory.

In some embodiments, the ankle joint connects the foot member to a firstend of the lower leg member of the apparatus and the knee joint isconnected to an opposite end of the lower leg member. In someembodiments, the apparatus includes the location is the ankle joint.

In some embodiments, the method includes adjusting the articulation ofthe foot member to a heel down position when the predetermined conditionis representative of the presence of level ground, an ascending ramp, ora descending ramp in underlying terrain. In some embodiments, the methodincludes adjusting the articulation of the foot member to a toe downposition when the predetermined condition is indicative of the presenceof an ascending stair or a descending stair in underlying terrain. Insome embodiments, the foot member is adjusted to a dorsiflexed positionrelative to a lower leg member of the apparatus when the predeterminedcondition is representative of the presence of an ascending stair. Insome embodiments, the foot member is adjusted to a plantar flexedposition relative to a lower leg member of the apparatus when thepredetermined condition is representative of the presence of adescending stair.

In some embodiments, the method includes adjusting the articulation ofthe foot member to a heel down position when the predetermined conditionis representative of the presence of level ground, an ascending ramp, ora descending ramp in underlying terrain, and adjusting the articulationof the foot member to a toe down position when the predeterminedcondition is representative of the presence of an ascending stair or adescending stair in underlying terrain. In some embodiments, thetrajectory is determined based on an inertial pose of a lower leg memberof the apparatus and an angle between the foot member and lower legmember. In some embodiments, the method includes adjusting thearticulation of the foot member of the apparatus to a predeterminedorientation when the trajectory satisfies a predetermined condition.

The invention, in another aspect, features an active lower extremityprosthetic, orthotic or exoskeleton apparatus. The apparatus includes afoot member, a lower leg member, and an ankle joint for connecting thefoot member to the lower leg member. The apparatus also includes a firstactuator for applying torque to the ankle joint to rotate the footmember with respect to the lower leg member and an inertial measurementunit for determining an inertial pose of the lower leg member. Theapparatus also includes a torque sensor for determining torque appliedto the lower leg member by the actuator. The apparatus also includes aforce sensor for determining axial force applied to the lower legmember. The apparatus also includes an angle sensor for determining anangle between the foot member and lower leg member. The apparatus alsoincludes a controller for controlling the actuator for modulating atleast one of joint impedance, position or torque of the apparatusthroughout a walking cycle of the apparatus based on the inertial pose,torque, axial force and angle.

In some embodiments, the apparatus includes one or more passive elasticmembers connected between the lower leg member and the foot member forstoring energy when the foot member rotates about the ankle joint towardthe lower leg member and for releasing energy to apply additional torqueto rotate the foot member away from the lower leg member. In someembodiments, the one or more passive elastic members is attached to theapparatus in parallel with the actuator. In some embodiments, the one ormore passive elastic members is a unidirectional spring and is notengaged during plantar flexion of the foot member relative to the lowerleg member.

In some embodiments, the actuator includes a series elastic actuator. Insome embodiments, the series elastic actuator comprises a brushlessmotor that drives a ball-screw, a carbon-fiber spring in series with anoutput of the ball-screw, and a strain sensor coupled to the spring. Insome embodiments, the inertial measurement unit comprises a three-axisrate gyro and a three-axis accelerometer.

In some embodiments, the apparatus includes a structural element coupledto the lower leg member and which also includes an interface forcoupling to a limb socket member of a wearer, wherein the structuralelement comprises a plurality of strain gages for determining the torqueapplied to the lower leg member by the actuator and the axial forceapplied to the lower leg member. In some embodiments, the actuatoradjusts stiffness of the apparatus during controlled plantar flexionphase of the walking cycle to minimize forefoot collisions with anunderlying surface. In some embodiments, the actuator controls impedanceand torque on the ankle joint of the apparatus during a late stancephase of the walking cycle based on at least one of ambulation speed,terrain context or terrain texture. In some embodiments, the actuatormodulates impedance of the apparatus based on a ground reaction forceand zero moment pivot coordinates imparted by an underlying surface ontothe foot member, the inertial pose of the lower leg member, the torqueapplied to the lower leg member by the actuator, the axial force appliedto the lower leg member, and the angle between the foot member and lowerleg member.

In some embodiments, the actuator modulates the impedance of theapparatus as the wearer stands up from a seated position or sits downfrom a standing position based on the inertial pose of the lower legmember, the angle between the foot member and lower leg member, and theground reaction force and zero moment pivot coordinates.

In some embodiments, the apparatus is used to treat drop foot gait. Insome embodiments, the apparatus is used to treat a wearer havinganterior muscle weakness, posterior muscle weakness, or a combinationthereof.

In some embodiments, the apparatus includes a thigh member, a knee jointfor connecting the thigh member to the lower leg member and a secondactuator for applying torque to the knee joint to rotate the lower legmember with respect to the thigh member. The apparatus also includes asecond inertial measurement unit for determining an inertial pose of thethigh member and a second torque sensor for determining torque appliedto the thigh member by the second actuator. The apparatus also includesa second force sensor for determining axial force applied to the thighmember. The apparatus also includes a second angle sensor fordetermining an angle between the thigh member and lower leg member. Theapparatus also includes a controller that controls the first and secondactuator for modulating an impedance of the apparatus throughout awalking cycle of the apparatus based on the inertial pose, torque, axialforce and angle determined using the first and second devices.

In some embodiments, the apparatus includes a torso member and a hipjoint for connecting the torso member to the thigh member. The apparatusalso includes a third actuator for applying torque to the hip joint torotate the thigh member with respect to the torso member. The apparatusalso includes a third inertial measurement unit for determining aninertial pose of the torso member and a third torque sensor fordetermining torque applied to the torso member by the third actuator.The apparatus also includes a third force sensor for determining axialforce applied to the torso member and a third angle sensor fordetermining an angle between the torso member and the thigh member,wherein the controller controls the first, second and third actuator formodulating an impedance of the apparatus throughout a walking cycle ofthe apparatus based on the inertial pose, torque, axial force and angledetermined using the first, second, and third devices.

In some embodiments, the lower leg member is attachable to a leg of thewearer. In some embodiments, the foot member is attachable to a foot ofthe wearer. In some embodiments, the thigh member is attachable to athigh of the wearer.

In some embodiments, the controller controls the actuator to modulate atleast two of joint impedance, position or torque of the apparatusthroughout a walking cycle of the apparatus. In some embodiments, thecontroller controls the actuator to modulate joint impedance, positionand torque of the apparatus throughout a walking cycle of the apparatus.

The invention, in another aspect, features a method for determining alevel of assistance desired for a lower-extremity orthotic orexoskeleton apparatus to apply to a wearer wearing the apparatus. Themethod includes specifying a physical therapy protocol defining a levelof assistance performed by the apparatus on the wearer over a period oftime and reducing the level of assistance performed by the apparatus onthe wearer to assist in rehabilitation of the limb pathology.

In some embodiments, the level of assistance by the apparatus is reducedbased on impedance and torque contribution of the wearer to theapparatus. In some embodiments, the apparatus includes the impedance andtorque contribution of the wearer is determined based on a biomechanicalmodel of the wearer and apparatus and measurements of the wearer duringoperation of the apparatus. In some embodiments, the measurements of thewearer include at least one of rotation and acceleration of at least onejoint of the apparatus. In some embodiments, the at least one joint ofthe apparatus includes at least one of a) an ankle joint that connects afoot member of the apparatus to a lower leg member of the apparatus; b)a knee joint for connecting a thigh member of the apparatus to the lowerleg member of the apparatus; or c) a hip joint for connecting a torsomember of the apparatus to the thigh member of the apparatus.

The invention, in another aspect, features a method for rehabilitationof a wearer with a limb pathology using a lower-extremity orthotic orexoskeleton apparatus worn by a wearer. The method includes estimatingimpedance and torque contribution of the wearer to at least one joint ofthe apparatus based on a biomechanical model of the wearer and apparatusand measurements of the wearer during operation of the apparatus andproviding a signal to an actuator of the apparatus that commands theactuator to provide additional torque to at least one joint of theapparatus such that a predetermined level of torque is achieved in theapparatus during operation.

In some embodiments, the measurements of the wearer include at least oneof rotation and acceleration of at least one joint of the apparatus. Insome embodiments, the at least one joint of the apparatus includes atleast one of a) an ankle joint that connects a foot member of theapparatus to a lower leg member of the apparatus; b) a knee joint forconnecting a thigh member of the apparatus to the lower leg member ofthe apparatus; or c) a hip joint for connecting a torso member of theapparatus to the thigh member of the apparatus.

The invention, in another aspect, features a method for estimating acondition of underlying terrain while a wearer is traversing theunderlying terrain. The method includes determining aninertially-referenced trajectory of points on a lower limb of a wearerand orientation of the lower limb of the wearer traversing underlyingterrain and analyzing the inertially-referenced trajectory relative toat least one predetermined trajectory model to estimate an underlyingterrain condition.

In some embodiments, the underlying terrain condition is at least one ofstair ascent, ramp ascent, level ground, ramp descent, or stair descent.In some embodiments, the method includes determining theinertially-referenced trajectory of the wearer traversing underlyingterrain, wherein the underlying terrain includes stair ascent, rampascent, level ground, ramp descent, and stair descent. In someembodiments, the determining an inertially-referenced trajectory of awearer traversing underlying terrain is performed during late swingphase of a gait cycle of the wearer.

In some embodiments, analyzing the inertially-referenced trajectoryrelative to at least one predetermined trajectory model includes usingat least one pattern recognition technique. In some embodiments, the atleast one pattern recognition technique is performed using a processorcoupled to at least one sensor and one actuator coupled to alower-extremity prosthetic, orthotic, or exoskeleton apparatus worn by awearer. In some embodiments, the at least one pattern recognitiontechnique is selected from the group techniques consisting of Bayesianpattern classification, neural nets, fuzzy logic or hierarchicaltemporal memory.

In some embodiments, the method includes controlling at least one ofjoint impedance, position or torque of a lower-extremity prosthetic,orthotic, or exoskeleton apparatus worn by a wearer based on theunderlying terrain condition estimate.

In some embodiments, the method includes determining a change intraction between a foot member of the apparatus and the underlyingsurface in which the apparatus includes a foot member, a lower legmember, an ankle joint for connecting the foot member to the lower legmember and a first actuator for applying torque to the ankle joint torotate the foot member with respect to the lower leg member. The methodcan include calculating ground reaction force imparted by the underlyingsurface onto the foot member based on an inertial pose of the lower legmember, the torque applied to the lower leg member by the actuator,axial force applied to the lower leg member, and an angle between thefoot member and lower leg member; and calculating velocity of the footmember zero moment pivot based on the inertial pose of the lower legmember, the torque applied to the lower leg member by the actuator, theaxial force applied to the lower leg member, and the angle between thefoot member and lower leg member.

In some embodiments, it is determined that the foot member is slippingor sinking if at least one component of the velocity of the foot memberzero moment pivot decreases during a portion of a gait cycle of thewearer between a foot-flat and toe-off condition. In some embodiments,the method includes reducing torque applied to the lower leg member inresponse to determining that the foot member is slipping or sinking. Insome embodiments, the method includes reducing the torque applied to thelower leg member by an attenuation factor. In some embodiments, theattenuation factor is a predetermined attenuation factor. In someembodiments, the attenuation factor is determined based on the zeromoment pivot velocity. In some embodiments, the method includes reducingthe torque applied to the lower leg member in response to the zeromoment pivot velocity being below a predetermined threshold.

The invention, in another aspect, features a method for discriminatingbetween properties of terrain underlying a lower extremity prosthetic,orthotic, or exoskeleton apparatus worn by a wearer, in which theapparatus includes a foot member, a lower leg member, and an ankle jointfor connecting the foot member to the lower leg member. The methodincludes estimating an inertial velocity vector attack angle of theankle joint of the apparatus throughout a gait cycle and discriminatingbetween terrain properties based on whether the inertial velocity vectorattack angle lies within a predetermined range.

In some embodiments, the method includes adjusting the impedance of theapparatus to minimize a cost function based on projected force impartedon the lower leg member during a period of time between when a heel ofthe foot member strikes the underlying terrain to when the foot memberis positioned in a flat-foot position relative to the underlyingterrain. In some embodiments, the method includes controlling at leastone of ankle joint impedance, position or torque of the apparatus basedon whether the inertial velocity vector attack angle lies within apredetermined range. In some embodiments, the foot member is attachableto a foot of the wearer and the lower leg member is attachable to a legof the wearer. In some embodiments, the foot member and lower leg memberreplace the foot and lower leg of the wearer.

The invention, in another aspect, features a method for controlling atleast one of joint impedance, position or torque of a lower extremityprosthetic, orthotic, or exoskeleton apparatus worn by a wearer, whereinthe apparatus includes a foot member, a lower leg member, and an anklejoint for connecting the foot member to the lower leg member. In someembodiments, the method includes estimating an inertial velocity vectorattack angle of the ankle joint of the apparatus throughout a gait cycleand adjusting the position of a foot member of the apparatus to a toedown position when the velocity vector attack angle lies within apredefined range.

In some embodiments, the method includes adjusting the position of thefoot member to a heel down position when the inertial velocity vectorattack angle is outside of the predetermined range. In some embodiments,the method includes adjusting the impedance of the apparatus to minimizea cost function based on projected force imparted on the lower legmember during a period of time between when a heel of the foot memberstrikes the underlying terrain to when the foot member is positioned ina flat-foot position relative to the underlying terrain. In someembodiments, the foot member is attachable to a foot of the wearer andthe lower leg member is attachable to a leg of the wearer. In someembodiments, the foot member and lower leg member replace the foot andlower leg of the wearer.

The invention, in another aspect, features a method of operating alower-extremity prosthesis or orthosis apparatus, in which the apparatusincludes a foot member and an ankle joint. The method includes trackinga trajectory of a portion of the apparatus and determining whether thetracked trajectory corresponds to stairs. The method also includesoptimizing operation of the apparatus for locomotion on stairs, insituations where the tracked trajectory corresponds to stairs. Themethod also includes determining whether the tracked trajectorycorresponds to non-stair terrain. The method also includes optimizingoperation of the apparatus for locomotion on non-stair terrain, insituations where the tracked trajectory corresponds to non-stairterrain.

In some embodiments, determining whether the tracked trajectorycorresponds to stairs comprises determining that a velocity vectorattack angle Ψ of the ankle joint in a late swing phase is below athreshold value, and the step of determining whether the trackedtrajectory corresponds to non-stair terrain comprises determining that avelocity vector attack angle Ψ of the ankle joint is above the thresholdvalue. In some embodiments, optimizing operation of the apparatus forwalking on stairs comprises adjusting a position of the foot member to atoe down position prior to foot strike, and wherein the step ofoptimizing operation of the apparatus for locomotion on non-stairterrain comprises adjusting a position of the foot member to a heel downposition prior to foot strike.

In some embodiments, optimizing operation of the apparatus for walkingon non-stair terrain includes dynamically controlling an impedance ofthe ankle joint during different phases of a single step, dynamicallycontrolling a position of the ankle joint during different phases of asingle step, and dynamically controlling torque of the ankle jointduring different phases of a single step.

In some embodiments, optimizing operation of the apparatus for walkingon stairs includes dynamically controlling an impedance of the anklejoint during different phases of a single step, dynamically controllinga position of the ankle joint during different phases of a single step,and dynamically controlling torque of the ankle joint during differentphases of a single step.

In some embodiments, the method includes determining whether the trackedtrajectory corresponds to an ascending ramp; optimizing operation of theapparatus for ascending a ramp in situations where the trackedtrajectory corresponds to an ascending ramp; determining whether thetracked trajectory corresponds to a descending ramp; and optimizingoperation of the apparatus for descending a ramp in situations where thetracked trajectory corresponds to a descending ramp.

In some embodiments, optimizing operation of the apparatus for ascendinga ramp includes dynamically controlling an impedance of the ankle jointduring different phases of a single step, dynamically controlling aposition of the ankle joint during different phases of a single step,and dynamically controlling torque of the ankle joint during differentphases of a single step, and the step of optimizing operation of theapparatus for descending a ramp includes dynamically controlling animpedance of the ankle joint during different phases of a single step,dynamically controlling a position of the ankle joint during differentphases of a single step, and dynamically controlling torque of the anklejoint during different phases of a single step.

In some embodiments, determining whether the tracked trajectorycorresponds to stairs includes determining that a velocity vector attackangle Ψ of the ankle joint in a late swing phase is below a thresholdvalue, and the step of determining whether the tracked trajectorycorresponds to non-stair terrain comprises determining that a velocityvector attack angle Ψ of the ankle joint is above the threshold value;wherein the step of optimizing operation of the apparatus for walking onstairs comprises adjusting a position of the foot member to a toe downposition prior to foot strike, and wherein the step of optimizingoperation of the apparatus for locomotion on non-stair terrain includesadjusting a position of the foot member to a heel down position prior tofoot strike, and wherein the step of optimizing operation of theapparatus for walking on non-stair terrain includes the steps ofdynamically controlling an impedance of the ankle joint during differentphases of a single step, dynamically controlling a position of the anklejoint during different phases of a single step, and dynamicallycontrolling torque of the ankle joint during different phases of asingle step.

The invention, in another aspect, features a lower-extremity prosthesisor orthosis apparatus that includes a foot member, a lower leg memberand an ankle joint operatively connected between the foot member and thelower member to permit articulation of the foot member with respect tothe lower leg member. The apparatus includes a motor configured to drivethe ankle joint and an inertial measurement unit configured to track atrajectory of the lower leg member and generate an output thatrepresents the trajectory. The apparatus also includes a controller thatis configured to (a) determine whether the tracked trajectorycorresponds to stairs based on the output, (b) optimize operation of theankle joint for walking on stairs when the tracked trajectorycorresponds to stairs, (c) determine whether the tracked trajectorycorresponds to non-stair terrain, and (d) optimize operation of theankle joint for walking on non-stair terrain when the tracked trajectorycorresponds to non-stair terrain.

In some embodiments, the controller determines whether the trackedtrajectory corresponds to stairs by determining that a velocity vectorattack angle Ψ of the ankle joint in a late swing phase is below athreshold value, and wherein the controller determines whether thetracked trajectory corresponds to non-stair terrain by determining thata velocity vector attack angle Ψ of the ankle joint is above thethreshold value. In some embodiments, the controller optimizes operationof the ankle joint for walking on stairs by adjusting a position of thefoot member to a toe down position prior to foot strike, and wherein thecontroller optimizes operation of the ankle joint for locomotion onnon-stair terrain by adjusting a position of the foot member to a heeldown position prior to foot strike.

In some embodiments, the controller optimizes operation of the anklejoint for walking on non-stair terrain by dynamically controlling animpedance of the ankle joint during different phases of a single step,dynamically controlling a position of the ankle joint during differentphases of a single step, and dynamically controlling torque of the anklejoint during different phases of a single step. In some embodiments, thecontroller optimizes operation of the ankle joint for walking on stairsby dynamically controlling an impedance of the ankle joint duringdifferent phases of a single step, dynamically controlling a position ofthe ankle joint during different phases of a single step, anddynamically controlling torque of the ankle joint during differentphases of a single step.

In some embodiments, the controller is further configured to (e)determine, based in the output, whether the tracked trajectorycorresponds to an ascending ramp, (f) optimize operation of the anklejoint for walking on an ascending ramp when the tracked trajectorycorresponds to an ascending ramp, (g) determine whether the trackedtrajectory corresponds to a descending ramp, and (h) optimize operationof the ankle joint for walking on a descending ramp when the trackedtrajectory corresponds to a descending ramp.

In some embodiments, the controller optimizes operation of the anklejoint for ascending a ramp by dynamically controlling an impedance ofthe ankle joint during different phases of a single step, dynamicallycontrolling a position of the ankle joint during different phases of asingle step, and dynamically controlling torque of the ankle jointduring different phases of a single step, and wherein the controlleroptimizes operation of the ankle joint for descending a ramp bydynamically controlling an impedance of the ankle joint during differentphases of a single step, dynamically controlling a position of the anklejoint during different phases of a single step, and dynamicallycontrolling torque of the ankle joint during different phases of asingle step.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other objects, feature and advantages of theinvention, as well as the invention itself, will be more fullyunderstood from the following illustrative description, when readtogether with the accompanying drawings which are not necessarily toscale.

FIG. 1A is a schematic illustration of the different phases of awearer's gait cycle over level ground.

FIG. 1B is a schematic illustration of the different phases of awearer's gait cycle ascending stairs.

FIG. 1C is a schematic illustration of the different phases of awearer's gait cycle descending stairs.

FIG. 2A is a schematic illustration of a method for determining anklejoint, heel and toe trajectories of a prosthetic, orthotic, orexoskeleton apparatus, according to an illustrative embodiment of theinvention.

FIG. 2B is a plot of experimental data showing ankle joint accelerationduring walking.

FIG. 3 is a schematic illustration of a method for determining footslope (heel height), according to an illustrative embodiment of theinvention.

FIG. 4 is a schematic illustration of a method for determining thecoordinates of the heel and toe in relation to the ankle joint in thefoot frame of reference, according to an illustrative embodiment of theinvention. The following is a reproduction of text previously presentedin FIG. 4.

-   -   Step 1: Establish ground reference (X) relative to ankle joint.    -   Step 2: Place toe on ground reference (X). Measure ankle joint        translation and foot rotation.    -   Step 3: Place heel on ground reference (X). Measure ankle joint        translation and foot rotation.    -   Step 4: Use simple trigonometry to determine the heel and toe        coordinates in the foot-frame.

FIG. 5 is a schematic illustration of a method for estimating a heelvector, according to an illustrative embodiment of the invention.

FIG. 6A illustrates the inertial measurement unit-computed ankle jointpivot trajectories in different ambulation contexts.

FIG. 6B illustrates the 2-D geometry that describes the in-flighttrajectory of the ankle joint of the prosthetic apparatus.

FIG. 6C illustrates how a stair-ramp discriminator can be built usingthe ankle angle attack angle as the trajectory feature thatdiscriminates between the stair and ramp ambulation context, accordingto an illustrative embodiment of the invention.

FIG. 7A illustrates a method for positioning an ankle joint prior tofoot strike, according to an illustrative embodiment of the invention.The following is a reproduction of text previously presented in FIG. 7A.

Algorithm

-   -   1. Wait until

${\begin{matrix}W_{k - 1} \\A\end{matrix}\overset{\rightarrow}{V}} < 0.$

-   -   2. Estimate f from the 6 DOF IMU state vector relative to ‘toe        off.’    -   3. Keep the ankle angle at θ* until heel or toe strike        (identified by series spring torque disturbance or other) where        θ* optimizes a cost criterion based upon the force of impact        (f_(ll)=M*V_(ll) ω_(N)) and the force of touchdown        (f_(σ)=M*V_(σ) ²/r//(γ+−φ) where r is the distance between the        ankle pivot and either the heel or toe.    -   4. Calculate the optimal complex impedance K*(s), necessary to        achieve the heel/toe down timing constraints.    -   5. Apply stance phase impedance control after heel/toe strike.

FIG. 7B illustrates how the method of FIG. 7A can be used to sense thepresence of stairs and overhang of the foot on the landing of the stair,according to an illustrative embodiment of the invention.

FIG. 7C illustrates a method for positioning an ankle joint in a rampambulation context, according to an illustrative embodiment of theinvention. The following is a reproduction of text previously presentedin FIG. 7C.

Algorithm

-   -   1. Wait until

${\begin{matrix}W \\A\end{matrix}{\overset{\rightarrow}{V}}_{Z}} < 0.$

-   -   2. Estimate φ from the 6 DOF IMU state vector relative to ‘toe        off.’    -   3. Keep the ankle angle at θ=        until heel or toe strike (identified by series spring torque        disturbance or other).    -   4. Apply impedance per stance phase control law at this        equilibrium point after heel/toe strike.

FIG. 7D illustrates how the method of FIG. 7B is adapted to use theoptimized impedance, according to an illustrative embodiment of theinvention. The following is a reproduction of text previously presentedin FIG. 7D.

Algorithm

-   -   1. Wait until

${\begin{matrix}W \\A\end{matrix}{\overset{\rightarrow}{V}}_{Z}} < 0.$

-   -   2. Estimate φ from the 6 DOF IMU state vector relative to ‘toe        off’    -   3. Keep the ankle angle at θ=φIMU until heel or toe strike        identified by series spring torque disturbance or other).    -   4. At θ=φIMU, calculate the complex impedance, K(s), that        satisfies the force of touchdown timing constraint.    -   5. Apply this impedance after heel/toe strike.

FIG. 8 illustrates a method for determining the inertially-referencedspring equilibrium based on the terrain angle at foot-flat.

FIG. 9 illustrates the effect of walking speed on ankle torque versusankle angle and shows how a push-pull actuator control applies to anappropriately selected parallel elastic element.

FIG. 10A illustrates a method for controlling a lower-extremityapparatus, according to an illustrative embodiment of the invention.

FIG. 10B is a schematic illustration of a model-based controller forimplementing impedance and torque control in a lower-extremityprosthetic apparatus, according to an illustrative embodiment of theinvention.

FIG. 10C is a schematic illustration of a model-based controller forimplementing torque control in a lower-extremity prosthetic apparatus,according to an illustrative embodiment of the invention.

FIG. 10D is a schematic illustration of the mechanical impedancerelation that governs the impedance control performed in FIG. 10A.

FIG. 10E is a schematic illustration of the impedance and reflexrelation that governs the impedance and reflex control performed in FIG.10B.

FIG. 10F is a schematic illustration of how zero moment pivot referencedground reaction forces are used to determine the restoring torquenecessary to stabilize inverted pendulum dynamics of a person wearing aprosthetic apparatus. The following is a reproduction of text previouslypresented in FIG. 10F.

f_(l) and f_(t) are the ground reaction forces acting on the leading andtrailing feet respectively. V_(CM) is the velocity vector of the wearercenter-of-mass. ZMP_(l) and ZMP_(t) denote the zero moment pivot on theleading and trailing feet. ^(w)r_(ZMP) _(l) ^(W)r_(ZMP) _(t) denote theworld coordinate referenced vectors between the center-of-mass and thezero-moment pivot on the leading and trailing feet respectively.

FIG. 11A is a schematic illustration of a lower leg foot member, anklejoint, and foot member of an ankle prosthesis showing ground reactionforces and the zero moment pivot.

FIGS. 11B-11D are schematic illustrations of the components of an ankleprosthesis showing the force and moment relationships among thecomponents necessary to determine the ground reaction forces and thezero moment pivot.

FIGS. 12A-12B illustrate the biomimetic (Γ-θ) behavior of an ankleprosthesis on level ground as a function of walking speed during poweredplantarflexion.

FIG. 12C-12D illustrate the effect of foot transitions on ground contactlength.

FIG. 12E illustrates how velocity-dependent tables of length of contactattenuation can use normalized ground contact length as a means toachieve biomimetic behavior during powered plantarflexion.

FIG. 12F illustrates how the estimated, y-component of the zero momentpivot vector changes during a typical walking motion.

FIG. 12G illustrates a method for incorporating an attenuation factorinto performance of an apparatus, according to an illustrativeembodiment of the invention.

FIG. 13A is a schematic representation of a control system scheme for aheel strike case, according to an illustrative embodiment of theinvention.

FIG. 13B is a schematic representation of a control system scheme for atoe strike case, according to an illustrative embodiment of theinvention.

FIG. 13C illustrates a method for position control applied to an ankleprosthesis (e.g., apparatus 1700 of FIG. 17A), according to anillustrative embodiment of the invention.

FIG. 14A illustrates a method for employing step-by-step terrainadaptation, according to an illustrative embodiment of the invention.The following is a reproduction of text previously presented in FIG.14A-A terrain adaptive robotic ankle prosthesis—Sensor processing andcontrol sequence optimizes ankle position, impedance and reflex for anyambulation task context “on-the-fly.”

Key

-   -   0: Ankle position during the last stance phase    -   0-1: Calculate ankle trajectory via numerical integration of the        IMU rate and acceleration    -   1-2: Determine the ambulation task type (level-ground, ramp,        stair, etc.) via intelligent modeling    -   3-4: Apply optimal “pre-strike” ankle position and impedance via        biomimetic modeling    -   4-5: Apply optimal impedance to eliminate foot slap    -   5-6: Apply optimal impedance and reflex to achieve biomimetic        optimization of net work and natural feel

FIG. 14B illustrates exemplary impedance that an ankle joint prosthesiswould apply for three different ambulation contexts.

FIG. 15 is a schematic representation of a lower-extremity biomechanicalsystem, according to an illustrative embodiment of the invention.

FIG. 16 illustrates a method of pose reconstruction for torso pose,thigh pose and torso/body center-of-mass pose, according to anillustrative embodiment of the invention.

FIG. 17A is an illustration of a lower-extremity prosthetic apparatus,according to an illustrative embodiment of the invention.

FIG. 17B is an illustration of a portion of the lower extremityapparatus of FIG. 17A that depicts a passive parallel elastic element.

FIG. 17C is an illustration of the passive parallel elastic element ofthe apparatus of FIG. 17B.

FIG. 17D is an illustration of the free-body diagram for the passiveparallel elastic element of FIG. 17C, according to an illustrativeembodiment of the invention.

FIG. 17E is an illustration of a perspective view of the structuralelement (pyramid) of the apparatus of FIG. 17A, according to anillustrative embodiment of the invention. The following is areproduction of text previously presented in FIG. 17E.

Regions of high strain. When strains in these two regions aredifferenced, a high-quality signal proportional to the moment load isobtained. Note: the difference signal is not sensitivity to axial load.

FIG. 17F is an illustration of a cross-sectional view of an alternativemethod for measuring axial force and moment applied to the lower legmember of FIG. 17A, according to an illustrative embodiment of theinvention.

FIG. 17G is an illustration of a method for computing the in-planemoment vector and axial force using a circular array of displacementsensors on a printed circuit assembly, according to an illustrativeembodiment of the invention.

FIG. 17H is a schematic illustration of a state and actuator controllerfor use with the apparatus of FIGS. 17A-17G, according to anillustrative embodiment of the invention.

FIG. 17I is a schematic illustration of an electrical circuit equivalentof a lower extremity prosthetic apparatus, according to an illustrativeembodiment of the invention.

FIG. 17J is a schematic illustration of the electrical circuit of FIG.17I including sensor measurements used in controlling the apparatus.

FIGS. 18A-18D are illustrations of a passive series-elastic member,according to an illustrative embodiment of the invention.

FIG. 19A is an illustration of a lower-extremity prosthetic apparatusincorporating a flat series spring, according to an illustrativeembodiment of the invention.

FIGS. 19B-19C are illustrations of a prosthetic apparatus using analternative series spring, according to an illustrative embodiment ofthe invention.

FIG. 20A is an illustration of a perspective view of a linear actuatorcapable of being used in various lower-extremity prosthetic, orthotic,and exoskeleton apparatus, according to an illustrative embodiment ofthe invention.

FIG. 20B is an illustration of a cross-sectional view of the linearactuator of FIG. 20A.

FIG. 21 is an illustration of a perspective view of a linear actuatorcapable of being used in various lower-extremity prosthetic, orthotic,and exoskeleton apparatus, according to an illustrative embodiment ofthe invention.

FIG. 22A is a schematic illustration of a top view of a lower-extremityorthotic or exoskeleton apparatus (wearable robotic knee brace),according to an illustrative embodiment of the invention.

FIG. 22B is a side view of the apparatus of FIG. 22A.

FIG. 22C is a schematic illustration of the interior portion of the kneejoint drive assembly of the apparatus of FIGS. 22A and 22B.

FIG. 23A is a schematic illustration of the human balance problem on aninclined slope.

FIG. 23B is a schematic illustration of acceptable solutions to thebalance problem based on variable knee flex by a wearer.

FIG. 23C is a schematic illustration representing the human body and howintrinsic sensing can be used to balance the wearer on level ground.

FIGS. 24A-24C are schematic illustrations for a method for balancing awearer as the wearer stands up from a chair, according to anillustrative embodiment of the invention.

FIG. 25A illustrates a definition of transfer work. The following is areproduction of text previously presented in FIG. 25A.

W_(t) is the transition work

f_(l) and f_(t) are the ground reaction forces acting on the leading andtrailing feet respectively

v_(cm) is the velocity vector of the wearer center-of-mass

T is the time from leading leg heel-strike to lift-off of the trailingleg

T is also referred to as the time during which the body isdouble-supported

1. Foot-strike impedance: Adjust late swing impedance in forward leg(spring equilibrium and dynamic stiffness) to minimize impact forceeffect on transfer energy and hip impact force/force rate

2. Torque: Assert reflex response in trailing leg to achieve positivevertical CG momentum in advance of the leading leg foot strike. Thiswill minimize negative contribution of leading leg impact and includethe positive contribution from trailing leg traction force —togetherminimizing the total transition work.

3. Inertially-referenced mid-stance impedance: To improve balance on thesloping terrain, apply inertially referenced restoring force inmid-stance.

FIG. 25B illustrates a definition of hip impact forces. The following isa reproduction of text previously presented in FIG. 25B.

W_(Û) _(l) and W_(Û) _(t) are unit vectors in world coordinates that liealong the leading and trailing thigh bones respectively.

f_(l) and f_(t) are the ground reaction forces acting on the leading andtrailing feet respectively.

f_(l) ^(hip) and f_(t) ^(hip) are the impact forces on the leading andtrailing hip joints respectively.

FIG. 26: Heel-Strike on Arbitrary Terrain with slope angle, φ.

FIG. 27: Toe-Strike on Arbitrary Terrain with slope angle, φ.

In FIGS. 26-27, trigger events are illustrated as circles. The segmentsthat chain events 1-3 capture the Controlled Plantar Flexion state forthe heel-strike case and the Controlled Dorsiflexion I state for thetoe-strike case. The segments that chain events 3-5 capture theControlled Dorsiflexion and Controlled Dorsiflexion II states for theheel-strike and toe-strike respectively. And the segments that chainevents 5-8 capture the Powered Plantar Flexion state. The swing phase“segment” connects event 8 and 1.

The net work, Work, applied by the Ankle-Foot System is defined as theline integral along the Γ-θ trajectory as it traverses the1-2-3-4-5-6-7-8-1 chain:Work=−∫₁ ⁸ Γdθ−∫ ₈ ¹ ΓdθSince the ankle work performed by the shank is zero during the swingphase (the 8-1 segment) only the first term of the equation needs to beevaluated.

FIG. 28: Biomechanics Study of Level-ground walking at different speeds.

DETAILED DESCRIPTION

Determining Activity Being Performed

Inertial Pose and Trajectory Estimation

FIG. 2 is a schematic illustration of a method for determining anklejoint 200, heel 212 and toe 216 trajectories of a prosthetic, orthotic,or exoskeleton apparatus (for example, apparatus 1700 of FIG. 17A) basedon the inertial pose of a lower leg member 220 coupled to the anklejoint 200, and the angle between the lower leg member 220 and footmember 208. Pose is the position and orientation of a coordinate system.The apparatus 1700 includes an inertial measurement unit 204 coupled tothe lower leg member 220. The inertial measurement unit 204 includes athree-axis rate gyro for measuring angular rate and a three-axisaccelerometer for measuring acceleration. Placing the inertialmeasurement unit on the lower leg member 220 collocates the measurementof angular rate and acceleration for all three axes of the lower legmember 220. The inertial measurement unit 204 provides asix-degree-of-freedom estimate of the lower leg member 220 pose,inertial (world frame referenced) orientation and ankle-joint 200(center of rotation of the ankle-foot) location.

In some embodiments, the lower leg member 220 pose is used to computethe instantaneous location of the knee joint. By using knowledge of theankle joint 200 angle (θ) the instantaneous pose of the bottom of thefoot 208 can be computed, including location of the heel 212 and toe216. This information in turn can be used when the foot member 208 isflat to measure the terrain angle in the plane defined by the rotationalaxis of the ankle joint/foot member. Mounting the inertial measurementunit 204 on the lower leg member 220 has advantages over other potentiallocations. Unlike if it were mounted on the foot member 208, the lowerleg member 220 mounting protects against physical abuse and keeps itaway from water exposure. Further, it eliminates the cable tether thatwould otherwise be needed if it were on the foot member 208—therebyensuring mechanical and electrical integrity. Finally, the lower legmember 220 is centrally located within the kinematic chain of the hybridsystem (referring to FIG. 15), facilitating the computation of the thighand torso pose with a minimum of additional sensors.

The inertial measurement unit 204 is used to calculate the orientation,_(ankle) ^(w)O position, _(ankle) ^(w)p, and velocity, _(ankle) ^(w)v,of the lower-extremity prosthetic apparatus in a ground-referenced worldframe. _(ankle) ^(w)O may be represented by a quaternion or by a 3×3matrix of unit vectors that define the orientation of the x, y and zaxes of the ankle joint in relation to the world frame. The ankle joint200 coordinate frame is defined to be positioned at the center of theankle joint axis of rotation with its orientation tied to the lower legmember 220. From this central point, the position, velocity andacceleration can be computed. For points of interest in, for example,the foot (e.g., the heel 212 or toe 216), a foot member-to-ankle jointorientation transformation, _(foot) ^(ankle)O(θ) relation: is used toderive the position using the following relation:

$\begin{matrix}{{\,_{{point}\text{-}{of}\text{-}{interest}}^{w}p} = {{\,_{ankle}^{w}p} + {{{\,_{ankle}^{w}O}(\gamma)}_{foot}^{ankle}{O(\theta)}\left( {{}_{}^{}{}_{{point}\text{-}{of}\text{-}{interest}}^{}} \right)}}} & {{EQN}.\mspace{14mu} 1} \\{\mspace{79mu}{where}} & \; \\{\mspace{79mu}{{{\,_{foot}^{ankle}O}(\gamma)} = \begin{bmatrix}1 & 0 & 0 \\0 & {\cos(\gamma)} & {- {\sin(\gamma)}} \\0 & {\sin(\gamma)} & {\cos(\gamma)}\end{bmatrix}}} & {{EQN}.\mspace{14mu} 2}\end{matrix}$where γ is the inertial lower leg member angle, and

$\begin{matrix}{{{\,_{foot}^{ankle}O}(\theta)} = \begin{bmatrix}1 & 0 & 0 \\0 & {\cos(\theta)} & {- {\sin(\theta)}} \\0 & {\sin(\theta)} & {\cos(\theta)}\end{bmatrix}} & {{EQN}.\mspace{14mu} 3}\end{matrix}$where θ is the ankle joint angle.

In this embodiment, the inertial measurement unit 204, including thethree-axis accelerometer and three-axis rate gyro, is located on theforward face at the top of the lower leg member 220 (as shown in, forexample, FIG. 17A). It is necessary to remove the effect of scale, driftand cross-coupling on the world-frame orientation, velocity and positionestimates introduced by numerical integrations of the accelerometer andrate gyro signals

Zero-Velocity Update

Inertial navigation systems typically employ a zero-velocity update(ZVUP) periodically by averaging over an extended period of time,usually seconds to minutes. This placement of the inertial measurementunit is almost never stationary in the lower-extremity prostheticapparatus. However, the bottom of the foot is the only stationarylocation, and then only during the controlled dorsiflexion state of thegait cycle. An exemplary zero-velocity update method, which is notimpacted by this limitation, for use with various embodiments of theinvention is described further below.

To solve this problem, orientation, velocity and position integration ofankle joint is performed. After digitizing the inertial measurement unitacceleration, ^(IMU)a, the ankle joint acceleration (^(IMU)a_(ankle)) isderived with the following rigid body dynamic equation:^(IMU) a _(ankle)=^(IMU) a+ ^(IMU) {right arrow over (ω)}X ^(IMU) {rightarrow over (ω)}X _(ankle) ^(IMU) {right arrow over (r)}+{right arrowover ({dot over (ω)})}X _(ankle) ^(IMU) {dot over ({right arrow over(r)})}  EQN. 4where ^(IMU){right arrow over (ω)} and ^(IMU){dot over ({right arrowover (ω)})} are the vectors of angular rate and angular acceleration,respectively, in the inertial measurement unit frame and X denotes thecross-product.

The relationship is solved _(ankle) ^(w)O=_(IMU) ^(w)O similarly as inEQNS. 1-3 using standard strapdown inertial measurement unit integrationmethods, in accordance with the following relationships known to oneskilled in the art:ankle^(w){circumflex over (Φ)}=^(w){circumflex over (Ω)}(^(w){circumflexover (ω)})_(ankle) ^(w){circumflex over (Φ)}EQN. 5

_(ankle)=^(w) â _(ankle)−[0,0,g] ^(T)  EQN. 6^(w) {circumflex over (p)} _(ankle)=

_(ankle)  EQN. 7_(foot) ^(w){circumflex over (Φ)}=_(ankle) ^(w){circumflex over(Φ)}_(foot) ^(ankle){circumflex over (Φ)}=_(ankle) ^(w){circumflex over(Φ)}Rotation_(x)({circumflex over (Θ)})  EQN. 8

_(heel)=

_(ankle)+^(w){circumflex over (Ω)}(_(ankle) ^(w){circumflex over(Φ)}[{circumflex over ({dot over (Θ)})}00]^(T))^(w) r_(heel-ankle)  EQN. 9

_(toe)=

_(ankle)+^(w){circumflex over (Ω)}(_(ankle) ^(w){circumflex over(Φ)}[{circumflex over ({dot over (Θ)})}00]^(T))^(w) r _(tow-ankle)  EQN.10^(w) {circumflex over (p)} _(heel)=^(w) {circumflex over (p)}_(ankle)+^(w) r _(heel-ankle)  EQN. 11^(w) {circumflex over (p)} _(toe)=^(w) {circumflex over (p)}_(ankle)+^(w) r _(toe-ankle)  EQN. 12^(w) r _(heel-ankle)=_(foot) ^(w){circumflex over (Φ)}^(foot)(r _(heel)−r _(ankle))  EQN. 13^(w) r _(toe-ankle)=_(foot) ^(w){circumflex over (Φ)}^(foot)(r _(tow) −r_(ankle))  EQN. 14In equations 5-14 above, the matrix, {circumflex over (Φ)}, will be usedinterchangeably with the orientation matrix, _(IMU) ^(w)O.

The world frame-referenced ankle joint velocity and position are thenderived at a point in time after the time of the previous zero-velocityupdate (i^(th) zero-velocity update) based on the following:^(w) v _(ankle)(t)=∫_(ZVUP(i)) ^(t)(_(IMU) ^(w) O)^(IMU) a _(ankle)dt  EQN. 15^(w) p _(ankle)(t)=∫_(ZVUP(i)) ^(tω) v _(ankle) dt  EQN. 16where ^(w)p_(ankle)(t=ZVUP(i)) is reset to zero for all i.

Through experimentation, using logs of inertial measurement unit dataacquired from an exemplary lower-extremity prosthetic apparatus (e.g.,lower-extremity prosthetic apparatus 1700 of FIG. 17A), we determinedthat the inertial measurement unit-referenced accelerations weresufficiently quiet early (see FIG. 2B at approximately 50.75 seconds and50.9 seconds when the z-acceleration is equal to about 1 g(approximately 9.8 m/s²) in the controlled dorsiflexion state and thevariance of the z-acceleration is less than a predetermined value(<0.005 g²)—indicating a period in time where the lower leg member 220is rotating about a stationary ankle joint 200. In another embodiment ofthis technique, a suitable quiet period can be detected on some part ofthe foot. Knowledge of the acceleration, angular rate and angularacceleration of the ankle joint can be combined with the knowledge ofthe sensed ankle angle (angle between the foot member and the lower legmember), angle rate and angle acceleration to calculate the accelerationof any point on the foot. Some point on the bottom of the foot can oftenbe used to perform a zero velocity update on successive gait cycles.Once this velocity is known, the velocity of the ankle joint can becomputed a posteriori. This velocity (rather than zero) can be used as areference from which the zero velocity update can be performed.

In the lower-extremity prosthetic apparatus, a quiet period nearlyalways exists in the Controlled Dorsiflexion state, so a zero-velocityupdate may be performed for every step that the wearer takes. Duringeach zero-velocity update, the velocity error contribution from each ofthree terms are preferably evaluated—the tip, δθ_(x) of the world framez-axis about the x-axis (the vector aligned with the ankle joint axis ofrotation during the zero-velocity update on the previous step); thetilt, δθ_(y) of the world frame z-axis about the y-axis (a vectordefined as the cross-product of the world-frame vertical (opposing thegravity vector) and the world frame x-axis); and the inertialmeasurement unit scaling along the vertical axis, δg. The values ofthese terms are used to correct the computed pose, inertial orientationand previous computed poses and inertial orientations of the differentcomponents of the apparatus (e.g., the lower leg member 1712 of FIG.17A).

While performing the orientation, velocity and position integration, asensitivity matrix, M(t) is calculated that relates the velocity errorthat would be introduced by the vector of errors, α=[δθ_(x) δθ_(y)δg]^(T). M(t) based on the following relationship:

$\begin{matrix}{{M(t)} = {\frac{\partial}{\partial\alpha}\left( {{{}_{}^{}{}_{}^{}}a_{ankle}} \right)}} & {{EQN}.\mspace{14mu} 17}\end{matrix}$in which, M(t) is integrated numerically to generate the overallterminal velocity sensitivity, M*,M*=∫ _(ZVUP) _(i-1) ^(ZVUP) ^(i) M(t)dt  EQN. 18In some embodiments, the vector of errors is expanded to includeaccelerometer bias offsets if these errors are significant, therebyincreasing the number of columns in M(t) and in M*. In this case, M*⁻¹takes the form of the Penrose pseudo-inverse or, by an optimalinnovations gain, K*. K* can be computed using standard optimal linearfiltering methods. To one skilled in the art, other terms can beincluded or used without loss of generality.

At the zero-velocity update for step i, the value of α that would havegenerated the estimated non-zero ankle joint velocity,^(w)v_(ankle)(ZVUP_(i)), is determined based on:α=M* ^(−1w) v _(ankle)(ZVUP_(i))  EQN. 19where α is the innovations correction vector. Since the non-zerovelocity results in part from noise in the accelerometers and angularrate measurements, not all of the innovations correction (α) is applied.Instead, the correction is scaled by a filtering constant (fraction), k,depending on the magnitude of the noise. At this point, the neworientation matrix (_(ankle) ^(w)O) and gravity magnitude (g) aredetermined based on:_(ankle) ^(w) O(ZVUP_(i) ⁺)=O _(x)(−kα(1))O _(y)(−kα(2))_(ankle) ^(w)O(ZVUP_(i) ⁻)  EQN. 20g(ZVUP⁺)=g(ZVUP⁻)−kα(3)  EQN. 21where O_(x)(tip) and O_(y)(tilt) denote incremental rotations of tip andtilt about the x and y axes respectively, and ZVUP_(i) ⁺ and ZVUP_(i) ⁻denote the times after and before the ZVUP, respectively.

It is possible to extend the zero velocity update to estimateaccelerometer and rate gyro bias offsets using linear estimators.Consistent angular alignment errors (e.g., about a given axis) could beused to estimate the rate gyro bias about that axis. In one embodiment,this is performed by creating linear stochastic models of accelerometerand rate gyro bias and using the zero-velocity update predictionresiduals as inputs to the linear filter applied to those models.

The above method is a method for continually updating the orientationand apparent gravity magnitude. An initialization procedure is used inthis embodiment when the lower-extremity prosthetic apparatus is firstpowered on. In this method, the wearer will, when requested by theapparatus (e.g., by a vibration code transmitted by the apparatus or analternative user interface), take one step forward and stop, then takeone step back to the original position. In this process, the steps willbe taken on the affected leg (for bilateral amputees, this calibrationwill be executed in a serial fashion as selected by the amputee). Thecalibration will invoke two ZVUP's—one to initialize the orientation andgravity magnitude, the second to check the result. This will ensureintegrity of the inertial measurement unit signals, processing andcontroller communication.

The above process accomplishes an initialization of the inertialorientation. It is, however, of general interest to accomplish a fullcalibration of the IMU, to account for the vector (ε) of error sources—avector that includes bias offset, scale and cross-sensitivity embodiedwithin the accelerometer and gyro signals. In manufacturing, a robot orother six degree-of-freedom machine can carry the IMU and applyreference trajectories in succession as a means of measuring the effectof these error sources. The sensitivity matrix (M(ε)) of the sensedreference trajectories to each of the error sources can be easilycomputed by those skilled in the art. By measuring the sensed deviationsfrom a rich set of reference trajectories—typically the deviation of theend-point of each trajectory segment—the vector (ε) can be estimatedusing regression or other linear estimation methods—provided that theset of reference trajectories is rich enough to excite the influence ofeach error source. The inventors have found that reference trajectoriesthat include closed paths like polygons and circles in three orthogonalplanes are sufficient to calibrate the full vector of error sources.Such reference trajectories can also be conducted by the wearer torecalibrate key elements of the vector (accelerometer bias, scale andcross-sensitivity) by, for example, walking in a sequence of closedpatterns on a horizontal plane and by rotating in sequence about avertical axis.

In some embodiments of the invention, these principles of the method aresimilarly applied to correcting or minimizing the effect ofaccelerometer and rate gyro drift error associated with accelerometersand rate gyros located on, for example, the thigh member and/or torso ofa wearer in which the prosthetic, orthotic, or exoskeleton apparatustreats or augments performance of these portions of a wearer's body. Inone embodiment, the method includes determining offset values for anaccelerometer signal and rate gyro signal output by an accelerometer andrate gyro coupled to a thigh member of the prosthesis or orthosis whenthe ankle joint is substantially stationary during a walking cycle ofthe prosthesis or orthosis. The method also can include measuring theangle of the lower leg member relative to the thigh member. In anotherembodiment, the method also includes determining offset values for anaccelerometer signal and rate gyro signal output by an accelerometer andrate gyro coupled to a wearer's torso when the ankle joint issubstantially stationary during a walking cycle of the prosthesis ororthosis. The method also can include measuring the angle of the thighmember relative to the wearer's torso. The methods can therefore beextended to the thigh member and/or torso of a wearer by performingthese measurements and relying on the linkage constraint relationshipsand related methods, as shown in FIG. 16. At the time of the zerovelocity update, the linkage constraints enable propagation of the jointvelocity references backwards from the ground-referenced zero velocityof the lowest link in the kinematic chain (e.g., the linkage thatdefines the hybrid human-robot system). These velocity references can beused as the input to the pose realignment and gravity compensation asdefined above.

Exemplary Ankle Joint Trajectories and Terrain Context Discrimination

Once the inertial measurement unit offsets have been calculated andcorrected (zeroed), the foot-slope (β) (alternatively referred to asheel height) is determined as illustrated in, for example, FIG. 3. Fromthe illustration it is easy to see that when the wearer is standing withher foot flat on the ground that β=−(θ+γ). By averaging over a period ofabout a tenth of a second an accurate estimate of β can be determined.Thereafter, the orientation component of the transformation that definesthe foot to ankle coordinate system, _(foot) ^(ankle)O, is computedbased on the following:

$\begin{matrix}{{{\,_{foot}^{ankle}O}\left( {\beta,\gamma} \right)} = \begin{bmatrix}1 & 0 & 0 \\0 & {\cos\left( {\beta + \gamma} \right)} & {- {\sin\left( {\beta + \gamma} \right)}} \\0 & {\sin\left( {\beta + \gamma} \right)} & {\cos\left( {\beta + \gamma} \right)}\end{bmatrix}} & {{EQN}.\mspace{14mu} 22}\end{matrix}$As before, the translational component of this transform will remainzero.

Once the foot-slope is defined, it is then necessary to determine theheel 212 and toe 216 coordinates in the foot coordinate system. In oneexemplary method for determining this, ^(foot){right arrow over(p)}_(heel) and ^(foot){right arrow over (p)}_(toc) are defined as thevector coordinates of the heel and toe in the new foot coordinatesystem. Because the rotational contribution of β has already beenincorporated, the z-component of these vectors is the same. It can beassumed that the x-component of these vectors are both zero. So thesevectors take the form:

$\begin{matrix}{{{}_{}^{}\left. p\rightarrow \right._{}^{}} = \begin{bmatrix}0 \\y_{heel} \\z_{0}\end{bmatrix}} & {{EQN}.\mspace{14mu} 23} \\{{{}_{}^{}\left. p\rightarrow \right._{}^{}} = \begin{bmatrix}0 \\y_{toe} \\z_{0}\end{bmatrix}} & {{EQN}.\mspace{14mu} 24}\end{matrix}$where z₀ defines the z-coordinate of the bottom of the foot (shoe).

FIG. 4 is a schematic illustration of a method for determining thecoordinates of the heel 212 and toe 216 in relation to the ankle joint200 in the foot frame of reference, according to an illustrativeembodiment of the invention. In the first step of the foot calibrationmethod defined in FIG. 4, the y-coordinate of the ankle joint 200 isaligned to a ground reference (e.g., seam in the pavement, a prominentfeature on a rug or on a linoleum surface). We arbitrarily define thisground reference to be the origin of the world coordinate system. Inmathematical notation, this alignment takes the form:

$\begin{matrix}{{{}_{}^{}{}_{{ankle}0}^{}} = \begin{bmatrix}0 \\0 \\{- z_{0}}\end{bmatrix}} & {{EQN}.\mspace{11mu} 25}\end{matrix}$where ^(world)p_(ankle) ₀ is the starting position for the moves thattake place in steps 2 and 3. In the second step, the toe 216 is placedonto the ground reference. In mathematical notation, this alignmenttakes the form:

$\begin{matrix}{\begin{bmatrix}0 \\0 \\0\end{bmatrix} = {{{}_{}^{}{}_{}^{}} + {{O(\gamma)}{{O\left( {\theta + \beta} \right)}\begin{bmatrix}0 \\y_{toe} \\z_{0}\end{bmatrix}}}}} & {{EQN}.\mspace{14mu} 26} \\{or} & \; \\{{{}_{}^{}\left. p\rightarrow \right._{}^{}} = {\begin{bmatrix}0 \\y_{toe} \\z_{0}\end{bmatrix} = {{- {O^{- 1}(\gamma)}}{O^{- 1}\left( {\theta + \beta} \right)}{{}_{}^{}{}_{{ankle}1}^{}}}}} & {{EQN}.\mspace{14mu} 27}\end{matrix}$

A similar relationship is determined during the alignment in step 3.When the equations above are solved independently, two differentestimates of z₀ are obtained. By combining the two constraint equationsinto one, a least-squares estimate of t_(heel), y_(toe) and z₀ can beobtained.

The heel 212 and toe 216 calibration method described above involves aseries of steps that would be used the first time a new pair offeet/shoes are worn. Such a calibration could be performed at, forexample, the prosthetist office.

In another exemplary method, the heel and toe vectors are calculatedon-the-fly. As shown in FIG. 5, the ankle joint 200 traces an arc 500 inthe early stance phase between foot-strike and foot-flat. The radius andorientation (midpoint angle) of the arc 500 fully determine the heel andtoe vectors. Mathematically, this is described as a series of anklepositions (^(world)p_(ankle) _(i) ) that are recorded during earlystance. Two ankle position measurements are needed, corresponding to twostatistically distinct lower leg member 220 (γ_(i′)) and ankle joint 200angle (θ_(i′)) positions, yielding:^(world) p _(heel) ₁ =^(world) p _(ankle) ₁ +O(γ₁)O(θ₁)^(foot) {rightarrow over (p)} _(heel)  EQN. 28^(world) p _(heel) ₂ =^(world) p _(ankle) ₂ +O(γ₂)O(θ₂)^(foot) {rightarrow over (p)} _(heel)  EQN. 29Then, by differencing the equations, the vector solution becomes:^(foot) {right arrow over (p)}_(heel)=(O(γ₂)O(θ₂)−O(γ₁)O(θ₁))⁻¹(^(world) p _(ankle) ₂ −^(world) p_(ankle) ₁ )  EQN. 30The solution requires that (O(γ₂)O(θ₂)−O(γ₁)O(θ₁)) is invertible. Andfrom an optimal linear filtering standpoint, this “gain matrix” must belarge enough so as to yield a statistically significant result.

Considering the fact that the lower-extremity prosthetic apparatusundergoes significant vibration during the early stance phase, theequations above can be extended to N sets of ankle joint position/anglemeasurements. The resulting N−1 equations can be solved usingleast-squares techniques to get an optimal estimate of the vector. Theequations above are similarly adapted to solve for the toe vector whentoe-strike initiates the early stance phase.

FIG. 6A illustrates the inertial measurement unit-computed ankle jointpivot trajectories in different ambulation contexts for a wearer walkingon various terrain: level ground (620), up a 5° ramp (624), down a 5°ramp (628), up a 10° ramp (632), down a 10° ramp (636), up stairs (640),and down stairs (644). Context is the shape of the terrain and how thewearer interacts with the terrain.

FIG. 6B illustrates the 2-D geometry that describes the in-flighttrajectory of the ankle joint of the prosthetic apparatus. If we treatlevel-ground walking as a subset of the ramp ascent/descent ambulationcontext (in which level ground is a zero degree ramp), then contextdiscrimination devolves into discrimination of stair ascent/descent fromramp ascent/descent. This discrimination is important because typicallyin the stair context, plantarflexion (rather than dorsiflexion) of theankle joint 600 is required to optimize foot-strike kinetics whereas inramp ambulation typically the ankle joint 600 is dorsiflexed (or heldneutral) to optimize foot-strike kinetics. In the latter context, it isonly in extremely steep descent that a plantar flexed ankle would be theappropriate orientation.

FIG. 6C illustrates how a stair-ramp discriminator can be built usingthe ankle angle attack angle (Ψ) as the trajectory feature thatdiscriminates between the stair and ramp ambulation context in a set ofrecorded data. FIG. 6C is a plot of the estimated velocity vector attackangle of the ankle joint 600 of the apparatus throughout a gait cycleversus each step taken by the wearer. In this data, an amputee fittedwith the prosthetic apparatus 1700 of FIG. 17A on his right foot tookthirty-one (31) steps (meaning walking cycles referenced to the rightfoot) in the following manner:

1. Steps 1-6: Six (6) steps up the 5° ramp

2. Step 7: One (1) step on the landing

3. Steps 8-9: Three (3) steps down the 10° ramp

4. Recording gap

5. Steps 10-11: Two (2) steps up the stairs

6. Step 12: One (1) step on the landing

7. Step 14-17: Four (4) steps down the 5° ramp

8. Steps (18-19): Two (2) steps on level-ground

9. Steps (20-21): Two (2) steps up the 10° ramp

10. Step (22): One (1) step from the 10° ramp to the landing

11. Steps (23-24): Two (2) steps down the stairs

12. Steps (25-31): Seven (7) steps on level-ground.

The steps taken during this recording included both ramp and stairascent and descent. FIG. 6C shows that stairs can be differentiated fromramps while the ankle is in-flight prior to foot-strike by monitoringthe ankle velocity attack angle (Ψ). When Ψ drops below a small positivevalue in this recording (and other similar recordings) the foot 604always lands on a stair. In all other cases, the foot lands on a ramp,irrespective of ramp angle (0°, −5°, +5°, −10°, +10°). Ψ is therefore asuitable ambulation task context discriminator that can be used by theprocessor to determine what activity is being performed.

Alternative methods for stair-ramp discrimination can be employed inother embodiments of the invention. The attitude (orientation ininertial space) lower leg member 608 (shank) and the ankle velocityattack angle (Ψ) can be used in one embodiment of the invention todistinguish between stairs or a ramp/level ground. The trajectory of theankle joint 600 in the y-z plane (referring to FIG. 6A) could be used inan alternative embodiment of the invention for stair-rampdiscrimination.

Swing Phase Ankle Positioning

The stair ramp discriminator provides a real-time prediction of theterrain slope angle,

. If the discriminator detects a step, including level-ground, then

=0. Otherwise, the slope angle is assumed to be:

$\begin{matrix}{\hat{\phi(t)} = {\min\left( {{\tan^{- 1}\left( \frac{\left( {{{{}_{}^{}{}_{}^{}}(t)} - {{{}_{}^{}{}_{}^{}}(0)}} \right)_{z}}{\left( {{{{}_{}^{}{}_{}^{}}(t)} - {{{}_{}^{}{}_{}^{}}(0)}} \right)_{y}} \right)},{\tan^{- 1}\frac{\left( {{{{}_{}^{}{}_{}^{}}(t)} - {{{}_{}^{}{}_{}^{}}(0)}} \right)_{z}}{\left( {{{{}_{}^{}{}_{}^{}}(t)} - {{{}_{}^{}{}_{}^{}}(0)}} \right)_{y}}}} \right)}} & {{EQN}.\mspace{14mu} 31}\end{matrix}$This slope angle corresponds to the minimum value possible given thatthe foot has not struck the ground.

is this the minimum value of two possible slope angles—the angle thatthe heel currently makes relative to the toe position from the last stepand the angle that the toe makes relative to the toe position from thelast step.

Once

is known, it is possible to apply various different methods to positionthe ankle in a way that adapts to this predicted terrain slope. Twoexamples of such methods are described below. In one embodiment of theinvention, the discriminator methodology described above is used tocontrol at least one of joint impedance, position or torque of a lowerextremity prosthetic, orthotic, or exoskeleton apparatus worn by awearer (e.g., the apparatus 1700 of FIG. 17A). The method involvesestimating a velocity vector attack angle of the ankle joint of theapparatus throughout a late swing (e.g., the y-axis values of the datain FIG. 6C). In one embodiment, the method also involves adjusting theposition of the foot member of the apparatus to a toe down position whenthe velocity vector attack angle has a predetermined sign (e.g., anegative value in the case of the data in FIG. 6C). In an anotherembodiment of the invention, the method involves adjusting the positionof the foot member of the apparatus to a heel down position when thevelocity vector attack angle has an opposite sign as the predeterminedsign (a positive sign).

In some embodiments, the method includes adjusting the impedance of theapparatus (e.g., the ankle joint impedance) to minimize a cost functionbased on projected force imparted on the lower leg member during aperiod of time between when a heel of the foot member strikes theunderlying terrain to when the foot member is positioned in a flat-footposition relative to the underlying terrain.

FIG. 7A illustrates a method for positioning the ankle joint 700 priorto foot strike. In this method, the ankle joint angle is optimized so asto minimize a cost functional based upon the projected force (f(t))imparted on the ankle joint 700 from foot member 708 strike tofoot-flat. Both heel-first 716 and toe-first 712 strategies areevaluated, and a strategy, including optimal ankle joint 700 angle,which minimizes the cost functional is selected. FIG. 7A describes themethod used.

In another embodiment, the method of FIG. 7A is augmented as shown inFIG. 7B to sense the presence of stairs, and to constrain theangle-of-attack optimization to toe-strike only in the event of stairswith short landing areas. For ascending or descending a steep, narrowset of stairs, the prosthetic apparatus is programmed to keep track ofthe volume swept by the foot during ascent—a volume for which there hasbeen no contact between the foot and the stairs. If in late swing, thereis determined to be no landing area for, for example, the heel, theoptimization is constrained to be the toe-down solution. In thisembodiment, a z-rotation is a rotation about the longitudinal axis ofthe lower leg member 704 (e.g., the z-axis of FIG. 17A) of theapparatus. If one descends stairs and rotates the foot member 708 inthis way, it is likely that the landing area is limited and the footmember 708 must be rotated to land squarely on the stair. In this case,the toe 712 down landing yields the only available minimum forcesolution for the method of FIG. 7A. Such z-rotation would signal thesystem that the landing area is limited, making a toe-down landing thesafest alternative when compared to heel-down.

The complex impedance computation employed in the method above can beapplied to any adaptive ankle positioning method as a means ofminimizing foot slap or use of excessive braking force as the anklejoint 700 rotates to the foot down state. FIG. 7D illustrates how themethod of FIG. 7A is adapted to use the optimized impedance. Once theoptimum angle-of-attack (

*) is found, an optimal control (Γ_(c)*(t)) is found that will bring thelinear and angular momentum of the ankle joint to zero withoutfoot-slap. The corresponding ankle angle response (θ*(t)) is then usedas the equilibrium trajectory. A corresponding optimal impedance, inrelation to this optimal trajectory, can be derived to accommodate theuncertainty in the momentum and the local terrain angle.

A simpler method can also be used as shown in FIG. 7C. FIG. 7Cillustrates a method for positioning the ankle joint in a rampambulation context. In this method, the ankle joint 700 angle isarticulated so as to be in the foot-flat position on a sloped-terrain(with slope angle

) when the lower leg member 704 is vertical. It is also useful togeneralize this method to adjust the ankle angle to be linearly relatedto the predicted slope angle by the relation:θ(t)=k

+θ ₀  EQN. 32Using this relationship the ankle angle can be adjusted to suit thewearer preferences.

In either of the two methods described above, the ankle joint angle 700prior to foot-strike will be controlled (steered) continuously tocoincide with the desired ankle joint 700 angle until the foot strikesthe ground.

Stance Phase Impedance and Torque Control

The next step involves restoring the orientation of the lower leg(shank) to align with the local vertical during stance phase. FIG. 8illustrates a method for determining the inertially-referenced springequilibrium based on the terrain angle at foot-flat of a lower-limbprosthesis 800, for example, the prosthesis apparatus 1700 of FIG. 17A.The prosthesis 800 has a foot member 808 with a toe 816 and heel 820.The prosthesis also has an ankle joint 804 and lower leg member (shank)812. The terrain angle (φ) is an input to the control system. Thecontrol system shifts the curve (Γ-Θ) (thereby altering the impedance ofthe ankle joint K_(controlled plantarflexion)) in FIG. 10A based on thechange in terrain angle (φ) to maintain or improve the overall balance(as described an illustrated in FIG. 10F) of the wearer duringcontrolled plantarflexion. The control system sets the impedance of theankle joint 804 of the prosthesis such that the ankle equilibrium angleis equal to the terrain angle (φ); and the control system restores theorientation of the lower leg member 812 (shank) to align with the localvertical 850.

FIG. 9 illustrates the effect of walking speed on ankle torque versusankle angle during controlled dorsiflexion. The control system shiftsthe curve (Γ-Θ) (thereby altering the impedance of the ankle joint 804K_(controlled dorsiflexion)) in FIG. 10A based on the change in terrainangle (φ) to maintain or improve the overall balance of the wearerduring controlled plantarflexion by commanding the ankle joint 804 tomove the lower leg member (shank) 812 towards the equilibrium point.

FIG. 10A illustrates a method for controlling a lower-extremityapparatus, according to an illustrative embodiment of the invention. Asshown in FIG. 10A, this is achieved in the control system by

-   -   1) adjusting the late swing impedance (step 1000) (the dynamic        stiffness and ankle-angle equilibrium angle) so as to soften the        impact between the time interval between foot-strike and        foot-flat, as described herein with respect to FIG. 7A (the        controller shifts the curve (Γ-Θ) (thereby altering the        impedance of the ankle joint K_(Powered plantarflexion)) based        on a minimization of the negative transfer energy impact and hip        impact forces during powered plantarflexion.    -   2) introducing a lifting force in the trailing leg—accomplished        by asserting a reflex response in the ankle (and knee) at or        before the time of impact of the leading leg (step 1004); and    -   3) maintaining an inertially-referenced equilibrium angle in the        controlled dorsiflexion phase to maintain balance (equilibrium)        (as described an illustrated in FIG. 10F) on sloping terrain        (step 1008).

FIG. 10B is a schematic illustration of controller for implementingimpedance and torque control in a lower-extremity prosthetic apparatus(e.g., the apparatus 1700 of FIGS. 17A-17E), according to anillustrative embodiment of the invention. FIG. 10E is a schematicillustration of the impedance and reflex relation that governs theimpedance and reflex control performed in FIG. 10B.

As shown, the spring, damping and inertial components of the impedanceare defined in relation to a trajectory, θ₀(t). Both the impedance gainmatrix and trajectory illustrated in FIG. 10B are loaded adaptively andin real-time from the state controller processor in accordance with thephase in the gait cycle, the terrain context, terrain texture andwalking speed as described above.

Studies have shown that intact limbs exhibit reflex responses that arisefrom non-linear positive torque (force) and non-linear positive jointvelocity feedback. The reflex relations as illustrated in FIG. 10Eemploy both types of feedback. Other non-linear implementations of thesepositive feedback relations can be used, including piece-wise linear andpiece-wise non-linear as would be obvious to those skilled in the art.In the preferred embodiment, positive torque feedback is achieved bymeasuring the torque in the shank of the ankle prosthesis and employingthis as the non-linear feedback signal, {circumflex over (Γ)}. In otherimplementations, this reflex torque input can be estimated using amodel-based computation of ankle dynamics.

The inventors have observed that the biomimetic impedance and reflex instance are coupled when the effects of walking speed and terrain slopeare taken into account as shown in FIG. 9. For this reason, in onepreferred embodiment, the parallel elasticity (e.g., parallel, or K3spring) for the prosthesis is picked so as to represent the stiffnessfor the slow walking speed as shown. In biomimetic systems, thestiffness component of the prosthesis is attenuated at higher walkingspeeds and the reflex response is steeper as shown in FIG. 9. Throughthis optimal biomimetic control and mechanical implementation, theresponse then requires the actuator to push on the parallel spring incontrolled dorsiflexion and to pull on it in powered plantar flexion. Wecall this bipolar, or push-pull, operation. In non-optimal control andmechanical implementations, the reflex is implemented by a unipolar,pulling-force-only of twice the magnitude. The preferred embodimentthereby reduces the peak actuator force and motor current by a factor oftwo, thereby extending the actuator design-life by 8× and reducingball-nut speed by nearly a factor of two when an appropriate bilateralseries spring response is chosen. This has tremendous advantages inincreasing the actuator durability, reducing actuator weight—the numberof ball-bearings and ball-nut diameter needed to achieve a design lifetarget are reduced—and reducing acoustic noise.

FIG. 10C is a schematic illustration of a controller for implementingimpedance control in a lower-extremity prosthetic apparatus (e.g., theapparatus 1700 of FIGS. 17A-17E), according to an illustrativeembodiment of the invention. FIG. 10D is a schematic illustration of themechanical impedance relation that governs the impedance controlperformed in FIG. 10C. τ_(M) is the torque applied by the linearactuator to the ankle joint of a lower extremity prosthetic apparatus.Through suitable “high-gain” compensation, G_(c)(z), where z denotes adiscrete-time signal transform, it is obvious that the motor torque willwork to make the sum of the torques applied by 1) the series-elasticactuator, 2) the “K3” parallel elasticity and 3) the acceleration torqueon the ankle equal to the torque command, Γ_(c), which is the desiredresult. The {circumflex over (K)}₃ and {circumflex over (K)}_(s) are areused to denote model estimates for these mechanical parameters, hencethe reference to model-based control.

FIG. 10F is a schematic illustration of how zero moment pivot referencedground reaction forces are used to determine the restoring torquenecessary to stabilize inverted pendulum dynamics of a person wearing aprosthetic apparatus. The torque (Γ_(CM)) is applied at thecenter-of-mass of the system (combination of, for example, the personwearing the prosthetic and the prosthetic) to maintain the balance ofthe wearer based on the following:Γ_(CM)=^(w) r _(ZMP) _(l) ×f _(l)+^(w) r _(AMP) _(t) ×f _(t)  EQN. 33where f_(l) and f_(t) are the ground reaction forces acting on theleading and trailing feet, respectively. v_(CM) is the velocity vectorof the wearer center-of-mass. ZMP_(l) and ZMP_(t) denote the zero momentpivot on the leading and trailing feet. ^(w)r_(ZMP) _(l) and ^(w)r_(ZMP)_(t) denote the world coordinate referenced vectors between thecenter-of-mass and the zero moment pivots on the leading and trailingfeet respectively. The term zero moment pivot refers to theinertially-referenced point on the foot about which the moment of theground reaction force distribution is zero. We will also refer to thispoint as the center-of-pressure (CoP) interchangeably throughout theremainder of this document.

Ground Reaction Forces and Zero Moment Pivot

Ground reaction forces (GRF) are the forces imparted by and underlyingsurface onto the foot (or foot member of a lower-extremity apparatus).Ground reaction forces are important biomechanical inputs during stance.By knowing the aggregate ground reaction force acting at the zero momentpivot (referred to as ZMP and CoP herein), the control system (e.g.,controller 1712 of FIG. 17A) of a lower-extremity prosthetic apparatushas a direct way of improving balance (of the wearer) and of optimizingpower delivery during the stance phase. U.S. Pat. No. 7,313,463, issuedto Herr et al. further describes estimating ground reaction forces andthe zero moment pivot position as well as biomimetic motion and balancecontrollers for use in prosthetics, orthotics and robotics and methods(the entire contents of which are hereby incorporated by reference inits entirety).

FIG. 11A is a schematic illustration of a lower leg foot member 1100,ankle joint 1104, and foot member 1108 of a prosthesis (e.g., apparatus1700 of FIG. 17A) that shows how the GRF components (specifically thevector from the ankle joint 1104 to the ZMP, ^(w)r_(ZMP), and the GRFvector, ^(w)F_(GRF)) change during the stance phase in a typical walkingcycle. The GRF estimation in research settings is often accomplished byapplying sensors on the sole of the shoe. But, such extrinsic sensingmay not be practical in prosthetic and orthotic devices because reliablepackaging means should preferably survive the contact stresses overmillions of walking cycles; which the sensors typically used in aresearch setting are unable to do so. Further, such means often requirecustomization of the shoe which is often not acceptable to the wearer.

In another embodiment of the invention, intrinsic sensing of the GRF isaccomplished in a novel way by combining inertial state and lower legmember force/torque inputs 1112 (using, for example, the structuralelement 1732 of FIGS. 17A and 17E).

FIGS. 11B, 11C, and 11D are schematic illustration of the components ofthe apparatus 1700 of FIG. 17A. The figures also show the force andmoment relationships among the components (linear series elasticactuator 1116 (e.g., combination of linear actuator 1716 plus serieselastic member 1724 of FIG. 17A) and parallel spring 1120 (e.g., passiveelastic member 1724 of FIG. 17A) necessary to determine the groundreaction forces and the zero moment pivot. ^(w)r_(ZMP) and ^(w)F_(GRF)are computed based on the following steps:

-   -   1. Update inertial state of the lower leg member 1100 and foot        member 1108 using inertial measurement unit and ankle joint 1104        angle inputs. Using rigid-body assumptions, further calculate        the world-referenced acceleration measured at the        center-of-mass (CM) of the lower leg member 1100 and foot member        1108 and the angular velocity and acceleration of the lower leg        member 1100 and foot member 1108.    -   2. Solve for F∥ as a function of the forces acting upon the        lower leg member 1100 as these are resolved along the lower leg        member 1100 axis.    -   3. Solve for F^(⊥) as a function of the moments applied by each        of the force and moment components acting upon the lower leg        member 1100.    -   4. Solve for ^(w)F_(GRF) using the values for F∥ and F^(⊥)        computed in steps 2 and 3 above and then balancing the forces        applied on the foot member 1100.    -   5. Balance the moments about the ankle joint 1104 assuming that        ^(w)F_(GRF) is applied at the foot-ground boundary (i.e.,        ^(w)r_(ZMP) ^(z)=0).    -   6. Solve for ^(w)r_(ZMP) ^(y).

Ankle Joint Behavior Due to Terrain Texture

FIG. 12A illustrates the biomimetic Γ-θ behavior of a prostheticapparatus (e.g., apparatus 1700 of FIG. 17A) on level ground as afunction of walking speed. FIG. 12B shows that the applied ankle jointtorque diminishes rapidly with angle during powered plantarflexion,thereby departing from the ideal biomimetic response and therebysignificantly reducing the net work performed (area under the Γ-θcurve), particularly when walking at high speed.

In conventional robotic systems, trajectories or other playback meansare employed to deliver repeatable and programmable responses. Suchmeans are not preferable in prosthetic and orthotic devices becausewearer intent may change in the middle of a playback segment. Forinstance, the wearer might be walking fast, then suddenly stop in frontof a patch of ice for instance. If pre-programmed trajectories or otherare played back, there is no easy way of aborting them without rapidchanges of force and torque—and without introducing hazards. Indeed,that is why the intrinsic means are used.

To extend the application of ankle joint torque during poweredplantarflexion, walking speed-dependent normalized ground contact lengthare used as the means of attenuating the peak plantarflexion torque, Γ₀.Ground contact length is estimated by using an idealized model of thefoot derived per the description related to FIGS. 2A-5 and by measuringthe inertial pose of the foot member during controlled dorsiflexion andpowered plantarflexion. As shown in FIG. 12C, as the foot transitionsfrom foot flat to toe-off, sections of the idealized foot will fallbelow the terrain, enabling an estimate of ground contact length. FIG.12D shows how L_(ground-contact) changes from foot flat to toe off.

FIG. 12E illustrates how velocity-dependent tables of Length of ContactAttenuation can use normalized ground contact length as a means toachieve biomimetic behavior during powered plantarflexion. The tablescan be computed by dynamically measuring the ground reaction force andfoot member pose of non-amputees in controlled environments as afunction of walking speed. The functional relationships between theattenuation function and ground contact length can be computed for eachwalking speed. These tables can be stored in the controller of theprosthetic apparatus as reference relationships. The functions can beshaped to suit specific wearer needs when the prosthetic apparatus isfitted to the wearer.

As described earlier, one of the motivations to use intrinsic feedbackas opposed to explicit trajectory or playback means is to accommodatechanges in wearer intent (e.g., decision to stop quickly). Intrinsicsensing using ground contact length as a means of attenuating anklejoint torque is not sufficiently general to accommodate changes inwearer intent involving stopping and changing direction. Referring toFIG. 12G, in one embodiment of the invention implemented on a prostheticapparatus, a time-dependent attenuation factor (e^(−t/τ)) is used inseries with the ground contact length attenuation. The time constant forthis attenuation, τ, can be picked so as to extinguish the poweredplantarflexion drive torque so as to prevent hazards associated withchanges in wearer intent. τ will typically range from 50-100 msec.

Preferably, the prosthetic apparatus enables the wearer to walk fasterwith less effort on all terrain. It is not sufficient to accommodatejust changes in terrain context (stairs, sloping ascent/descent).Changes in terrain texture as this might introduce slipping (e.g.,Ice/snow) or sinking (mud, snow, sand, fine gravel) hazards shouldpreferably be accommodated. Intrinsic sensing of the zero moment pivottrajectory can be used to optimize walking performance and/or toeliminate hazards while walking on varying terrain texture.

FIG. 12F illustrates how the estimated, y-component of the zero momentpivot vector, ^(w){circumflex over (r)}_(CoP) ^(y), changes during atypical walking motion. As shown, in a no-slip condition ^(w){circumflexover (r)}_(CoP) ^(y) must increase monotonically between the conditionsof foot-flat (3) and toe-off (4). This is because it is the heel that islifting off of the terrain surface during this period (increasingly asthe walking cycle progresses). If the velocity of the zero moment pivotever moves along the negative y-axis, the foot is slipping. In a fashionsimilar to how anti-lock brakes are implemented in vehicles, theprosthetic apparatus can reduce torque by an attenuation factor derivedfrom the integral of the negative zero moment pivot velocity. In oneembodiment, so as to reduce noise sensitivity, only negative velocitiesbelow a noise threshold are integrated.

FIGS. 13A and 13B provide the state control context for an illustrativeembodiment of the invention applied to, for example, apparatus 1700 ofFIGS. 17A-17E. Normal walking involves the cycling between two phases:the swing phase and the stance phase. FIG. 13A depicts a control systemscheme involving a walking motion in which the stance phase is initiatedby the heel striking 1320 the ground. _(a) ^(w)V_(z) denotes thez-component of the ankle joint velocity in the ground-referenced, worldframe. FIG. 13B shows a walking motion in which the stance phase isinitiated by the toe striking 1324 the ground.

Exemplary Control System Behavior for Driving Prosthesis or OrthosisThrough Gait Cycle

FIGS. 13A and 13B illustrate that the control system 1300 changes anklebehavior as the ankle transitions between states in the swing 1304 andstance phases 1308. The control system 1300 applies position control1328 in the swing phase—positioning the ankle so as to avoid triphazards in the early swing phase state and so as to optimize heel-toestrike attack angle (adaptive ankle positioning) for specific terrainconditions (slope, stairs, steps) in the late swing phase state. Thecontrol system 1300 applies impedance and torque control 1332 in thestance phase—optimizing the inertial, spring and damping characteristicsof the ankle—as the ankle transitions through the heel/toe strike, footdown, peak energy storage (dorsiflexion with exponential hardening),powered plantarflexion and toe-off events.

FIG. 13C illustrates a method for position control applied to a lowerlimb apparatus (e.g., apparatus 1700 of FIG. 17A), according to anillustrative embodiment of the invention. It is desirable to not movethe foot member 1348 forward until the wearer and/or the controller ofthe apparatus are sure the toe 1340 is going to clear the terrain infront of the wearer. One exemplary way to accomplish this is to waituntil the toe 1340 of the foot member 1348 is a sufficient distanceabove the last known position of the toe 1340 with respect to theunderlying terrain. In this embodiment, the control system 1300 appliesposition control 1328 by beginning to rotate the ankle joint 1340 onlyafter the clearance distance measured along a normal vector to theterrain surface between the toe 1340 of the foot member 1348 at time tand at time t_(k-1) is determined to be greater than (ε₀). Thisminimizes the risk that the toe 1340 will encounter a trip hazard. Inone embodiment, the position of the toe 1344 at the two different times(t and t_(k-1)) are determined using the inertial measurement unitmeasurements, as described previously herein. One skilled in the artwould understand how to apply other schemes to determine when it isappropriate to move the foot member 1348 forward. In some embodiments,the controller may determine it is appropriate to move forward based on,for example, whether the swept volume of the foot, when dorsiflexed,achieves the desired clearance relative to the terrain surface.

In summary, this embodiment of the invention, the prosthetic apparatusemploys step-by-step terrain adaptation with the intent to achieve truebiomimetic behavior in all ambulation task contexts; includinglevel-ground walking, stair ascent/descent and ramp ascent/descent. FIG.14A outlines the process by which the step-by-step adaptation isaccomplished. In the swing phase, the inertial measurement unit suppliesthe intrinsic sensing input (as opposed to say extrinsicneuronal/myoelectric inputs) that enables the apparatus to discernterrain context from cues supplied by swing-phase trajectory features.Adaptive swing-phase ankle positioning refers to the articulation of theankle angle, θ, so as to achieve a natural heel or toe touchdown that isoptimized for the most likely terrain context as determined by theterrain context discrimination on the swing phase trajectory cues.

FIG. 14B illustrates exemplary impedance that an ankle joint prosthesiswould apply for three different ambulation contexts. FIG. 14B is a graphof required ankle torque 1404 (units of Nm/kg) versus ankle joint angle1408 (units of degrees). The graph includes three curves 1412, 1416 and1420. Curve 1412 illustrates the ankle joint torque 1404 versus anklejoint angle 1408 for walking on a ramp incline of 5 degrees. Curve 1416illustrates the ankle joint torque 1404 versus ankle joint angle 1408for walking on a ramp decline of 5 degrees. Curve 1420 illustrates theankle joint torque 1404 versus ankle joint angle 1408 for walking on aramp incline of 0 degrees (level ground). The slope of the curves isequal to the stiffness (or impedance in general). The area enclosed bythe closed Γ-θ curve corresponds to the amount of non-conservative workrequired for the specific terrain context (e.g., slope, stairs) andwalking speed. As can be seen in the graphs, an ankle joint prosthesiswould be required to provide more work to accomplish the ambulation taskof walking up an inclined ramp versus walking on level ground becausethe area within the curve 1412 is greater than the area within the curve1416.

Generalization of the Hybrid Lower-Extremity Augmentation System

FIG. 15 is a schematic representation of a lower-extremity biomechanicalapparatus 1500, according to an illustrative embodiment of theinvention. In one embodiment, the apparatus 1500 is an orthoticapparatus that augments the ambulation abilities of the wearer. Inanother embodiment, the apparatus 1500 is an orthosis apparatus thatattaches to a wearer's body to support and/or correct musculoskeletaldeformities and/or abnormalities of a wearer's hip, thigh, lower leg andfoot. In another embodiment, the apparatus 1500 is an exoskeletonapparatus that attaches to a wearer's body to assist or augment thewearer's lower-extremity biomechanical output (e.g., augment thelower-extremity strength or mobility of the wearer).

The apparatus 1500 is a linkage represented by a plurality of links (ormembers) and joints that connect the links. The apparatus 1500 includesa foot member 1508 (L₀) coupled to a lower leg member 1516 (L₁) by anankle joint 1512. The apparatus 1500 also includes a thigh member 1524(L₂) coupled to the lower leg member 1516 by a knee joint 1520. Theapparatus also includes a hip joint 1528 that couples the thigh member1524 to the torso 1532 (L₃) of the wearer. Center-of-mass 1504 is thecenter-of-mass of the combination of the apparatus 1500 and the wearer.

The foot member 1508 contacts the terrain 1536 underlying the footmember 1508 at the zero moment pivot 1540. The foot member 1508 includesa toe portion 1544 and a heel portion 1548. Each joint of the apparatus1500 also includes an actuator with a generalized vector of torque(force) Γ_(i), displacement ξ_(i), and impedance K_(i), where i=0corresponds to the ankle joint 1512, i=1 corresponds to the knee joint,and i=2 corresponds to the hip joint. Each joint actuator may include amachine element (e.g., ball-screw actuator or rotary harmonic drive),human muscle, or both. Joint displacements typically take the form ofangular displacement (rotation) but may also include a combination oflinear and angular displacements as found in, for example, a typicalknee joint. The pose of a link, i, is represented by a 4×4 matrix thatdefines the location of the link origin and the unit vectors of itscoordinate frame in terms of the unit vectors in the world coordinateframe, W.

The pose of each link, j, can thereby be determined via linkageconstraint relationships—specifically by multiplying the pose of link,i−1, by a transformation defined by the generalized displacement, ξ_(j),and specific link parameters (link length, skew and convergence angles).For example, if the pose of the shank is known, the pose of the foot,thigh and torso can be computed provided that the generalizeddisplacements for those linkages are known, either by directly sensingthese or through the use of inertial sensors. The vector of sensorinformation intrinsic to each link is encapsulated in what we will callan intrinsic sensing unit (ISU). Examples of intrinsic sensors includedirect or indirect measurement of generalized displacements; measurementof the angular rate and acceleration of the link (e.g., using, forexample, an inertial measurement unit); measurement or estimation of thecomponents of force or torque on the link; multi-modal computer imagery(e.g., a range map) or measurement of the outputs of specific neuralpathways on or adjacent to the link.

The terrain is modeled as a contour function, z(x,y), with surfaceproperties, α(x,y). In this context, the surface properties wouldinclude the elasticity/plasticity, damping characteristics andcoefficient of friction of the surface sufficient to capture the abilityof the foot to gain traction on the surface and to capture the surfaceenergy as this would relate to the work required to touch down on thesurface and to push off of it with the foot member.

FIG. 16 is a schematic illustration of a method for determining the poseof the thigh member, hip member and torso of a wearer, according to anillustrative embodiment of the invention. In lower-limb systemsemploying robotic knee prostheses or orthosis, the location of the humanhip can also be computed, either by incorporating an inertialmeasurement unit on the thigh or by measuring the relative knee angle asreferenced to the lower leg member. If an inertial measurement unit isfurther employed on the torso, the pose of the torso can also beinstantaneously computed. Alternatively, the pose can be computed bymeasuring the two degree-of-freedom hip joint displacements.Compensation for the torso pose prediction errors arising from the rategyro and accelerometer drift on the torso inertial measurement unit canbe corrected during the lower leg member zero-velocity update through achaining of velocity constraints through the hybrid system linkages.

FIG. 16 illustrates a method of pose reconstruction in which j, j−1velocity constraints are used to correct the prediction of torso pose(_(torso) ^(w){circumflex over (T)}(t=t_(zvup) _(i) )), thigh pose(_(thigh) ^(w){circumflex over (T)}(t=t_(zvup) _(i) )) and torso/bodycenter-of-mass pose (^(w)v_(cg)(t=t_(zvup) _(i) )). Step 1 (1604)captures the output of the zero velocity update on the lower leg member1620 (link 1) to determine the lower leg member pose, as described abovewith respect to FIGS. 2A-5. The solutions (steps 2 and 3) for the thighmember 1624 (link 2) and torso member 1628 (link 3), respectively,follow the example of step 1 (1604), but in these cases the velocityconstraints are non-zero and are predicted by the translational androtational velocity from the prior link.

Exemplary Mechanical Designs

FIG. 17A is an illustration of a lower-extremity prosthetic apparatus1700, according to an illustrative embodiment of the invention. Theapparatus 1700 has a mounting interface 1704 making it capable ofattaching to a complementary lower-extremity limb socket member of awearer. The apparatus 1700 also includes a structural element 1732 (alsoreferred to herein as the pyramid) coupled to the mounting interface1704 and a first end 1752 of a lower leg member 1712 (also referred toherein as a shank). In some embodiments, the axial force and momentapplied to the lower leg member of the apparatus is determined based onsensor measurements made using the structural member (pyramid) coupledto the lower leg member of the apparatus. The pyramid is an instrumentedstructure that is a component of the prosthesis and which couples to thelimb socket of the wearer. In one embodiment, the pyramid (structuralelement) measurements are used by a controller to determine axial forceand moment applied to the lower leg member. In this embodiment, thestructural element 1732 is coupled to the first end 1752 of the lowerleg member 1712 with a set of pins 1711. The pins 1711 pass through aset of holes 1713 in the lower leg member 1712 and a set of holes 1715(shown in FIG. 17E) in the structural element 1732.

The structural element 1732 has a top surface 1731 located towards themounting interface 1704 and a bottom surface 1733 located towards thelower leg member 1712. The lower leg member 1712 is also coupled to afoot member 1708 at an ankle joint 1740 at a second end 1744 of thelower leg member 1712. The ankle joint 1740 (e.g., a rotary bearing)permits the foot member 1708 to rotate about the x-axis relative to thelower leg member 1712. The foot member includes a heel 1772 and a toe1776.

The apparatus 1700 also includes a linear actuator 1716 with a first end1736 and a second end 1748. The linear actuator 1716 generates a linearmotion 1703. The first end 1736 of the linear actuator 1716 is coupled(with, for example, a rotary bearing) to the first end 1752 of the lowerleg member 1712. The apparatus 1700 also includes a first passiveelastic member 1728 in series with the linear actuator 1716. The passiveelastic member 1728 is coupled to the foot member 1708 and the secondend 1748 of the linear actuator 1716. The passive elastic member 1728 iscoupled to the foot member 1708 (with, for example, a rotary bearing) atthe proximal end 1730 of the passive elastic member 1728. A distal end1726 of the passive elastic member 1728 is coupled between the secondend 1748 of the linear actuator 1716 (with, for example a rotarybearing). The linear actuator 1716 applies torque about the ankle joint1740.

The apparatus 1700 also includes an optional second passive elasticmember 1724 with a first end 1756 and a second end 1760. The secondpassive elastic member 1724 provides a unidirectional spring force inparallel (provides parallel elasticity) with the lower leg member 1712.The first end 1756 of the second passive elastic member 1724 is coupledto the first end 1752 of the lower leg member 1712. The second end 1760of the second passive elastic member 1724 is coupled to the foot member1708. However, during plantarflexion the spring is not engaged, andtherefore only provides a unidirectional spring force to the apparatus.

In some embodiments, the second passive elastic member 1724 is anon-compliant stop that stores little or no energy and limits furtherrotation of the ankle beyond a predefined angle during powered plantarflexion.

FIGS. 17B and 17C are illustrations of a portion of the lower extremityapparatus of FIG. 17A depicting the second passive elastic element 1724.The second passive elastic element 1724 stores energy duringdorsiflexion but, not in plantarflexion. The elastic element 1724 has adouble-cantilever engagement (clamped at a location 1780 between thefirst end 1756 and the second end 1760). The elastic member 1724 has atapered shape 1784 that causes the elastic member 1724 to provideefficient energy storage by maximizing bending strain along the entirelength (along the y-axis) of the elastic element 1724. In someembodiments, the normalized spring constant ranges from 0-12 Nm/rad/kg.At the high end of the range, the energy storage is approximately 0.25J/kg.

A cam/ramp arrangement of the elastic member 1724 facilitates tailoringof the spring constant to the weight of the wearer. The cam element 1788is located at the second end 1760 of the elastic member 1724. The rampelement 1792 is located on the foot member 1708. The cam element 1788engages the ramp element 1792 during dorsiflexion; however, the camelement 1788 does not engage the ramp element 1792 or another portion ofthe apparatus 1700 during plantarflexion. Because the cam element 1788does not engage the ramp element 1792 or another portion of theapparatus 1700 during plantarflexion, the elastic member 1724 storesenergy only during dorsiflexion. In one embodiment, the position of theramp element 1792 is screw-adjustable to allow the wearer or a secondparty to tailor the ramp engagement of the cam element 1788 so as toalign the energy storage characteristics to the wearer's walking habits.An operator may adjust the position of the ramp element 1792 relative tothe position of the cam element 1788 in order to modify the energystorage characteristics of the passive elastic member 1724.

In alternative embodiments, an actuator is integrated into the ramp toadjust the ankle joint angle at which the second passive elastic member1724 (elastic member engagement angle) engages. This would enable theankle joint 1740 to be dorsiflexed during the swing phase withoutengaging the elastic member 1724 when, for example, the wearer isascending ramps and stairs, and while running.

The passive elastic element 1724 also functions to increase thefrequency response of the apparatus 1700 when the elastic element 1724is engaged in dorsiflexion. The apparatus 1700 dynamics in dorsiflexionbenefit from a fast response (bandwidth) series elastic actuator (i.e.,combination of the linear actuator 1716 and first passive elasticelement 1728). The spring constant associated with the second passiveelastic element 1724 increases the bandwidth of the apparatus 1700 by afactor, β, where:β=(K ₃(1+K _(S) /K ₃)^(1/2) /K _(S))^(1/2)  EQN. 34where K₃ is the spring constant of the second passive elastic member1724 and K_(S) is the spring constant of the combination of the linearactuator 1716 and first passive elastic element 1728. In one embodimentof the invention, the second passive elastic element provides a β from 1to 3; thereby increasing the bandwidth of the apparatus 1700 from about5 Hz to about 15 Hz.

The second passive elastic member 1724 employs a dovetail feature 1796at both ends to enable clamping at both ends without use of mountingholes. In one embodiment, the second passive elastic member 1724 isfabricated from composite fiber materials. Mounting holes would form astress intensity and cause fiber dislocations in the passive elasticmember 1724 that would compromise the strength of the spring. The endclamps 1798 have complementary shapes that hold the passive elasticelement 1724 in place. In one embodiment of the invention, epoxy isemployed in the clamps to permanently secure the second passive elasticmember 1724 in the end clamps. The epoxy joint would be more prone tofailure in the absence of the dovetail features 1796.

The passive elastic element 1724 employs a tapered design to maximizeenergy storage in the element 1724 to ensure that energy storage densityis constant over its length for a given deflection. Referring to FIG.17D, we illustrate the free-body diagram for the passive elastic element1724, showing how the roller force, F_(roller), and the lower leg memberforce, F_(shank), combine to create an equal and opposite force by thecentral pivot. In this embodiment, the roller force and the lower legmember force are applied equidistant from the center pivot. The forcesat the end, F, combine to create a central pivot force of 17F. Usingstandard thin beam relationships, the moment acting at a distance of xfrom the central pivot varies linearly—starting at a value of FL in thecenter and falling to zero at x=L, where L is the length of the passiveelastic element 1724 between the locations at which the forces areapplied. Energy storage density along x is proportional to the productof moment (M(x)) and the strain at the surface (ε₀(x)), where:

$\begin{matrix}{{M(x)} = {F\left( {\frac{L}{2} - x} \right)}} & {{EQN}.\mspace{14mu} 35} \\{{\varepsilon_{0}(x)} = {\frac{M(x)}{{EI}\;\omega^{*}} = \frac{F\left( {\frac{L}{2} - x} \right)}{{EI}\;\omega^{*}}}} & {{EQN}.\mspace{14mu} 36}\end{matrix}$

For a given layup of composite material, the surface strain is keptbelow a critical value, ε*. For a given moment, the energy density inthe beam will be maximized when the surface strain is set to thiscritical value. To keep the energy density constant and at its maximumvalue, the optimal width of the beam, w*(x), is defined by the relation:

$\begin{matrix}{w^{*} = {{w^{*}(x)} = {\frac{F}{{EI}\;\varepsilon_{0}^{*}}\left( {\frac{L}{2} - x} \right)}}} & {{EQN}.\mspace{14mu} 37}\end{matrix}$

In one embodiment, the taper 1784 varies linearly from the center of thebeam. By using this design method, we have amplified the energy storageof the spring by over a factor of 2 when compared to a beam without ataper 1784. Because the composite spring material is not homogeneous andsince the thin beam equations are not applicable, computational toolsare used to estimate the energy storage density in the passive elasticmember 1724. The shape that is able to store the most energy is highlydependent upon the fiber laminate, lamination design, thickness and theexact manner in which the passive elastic member 1724 is attached to theapparatus 1700. We have determined, however, that a linear taperdelivers energy storage within about 10% of the optimal. In a preferredembodiment, the linear taper is used because of the relative ease bywhich a linear taper pattern maybe cut from a sheet of laminated plycomposite material using a water jet process. In alternative, lesspreferred embodiments, a non-tapered spring may be used.

FIG. 17E is an illustration of a perspective view of an embodiment ofthe structural element 1732 (also referred to herein as the pyramid).The structural element 1732 is coupled between the mounting interface1704 and the first end 1752 of the lower leg member 1712. The structuralelement 1732 is coupled to the first end 1752 of the lower leg member1712 with a set of pins 1711 (shown in FIG. 17A). The pins 1711 passthrough the set of holes 1713 in the lower leg member 1712 and a set ofholes 1715 in the structural element 1732. The pins 1711 allow a rotarydegree of freedom for the strain in structural element 1732 from beingfalsely recorded as axial force and moment in the structural element1732. In this embodiment, the structural element 1732 is capable ofmeasuring the moment and axial load on the ankle joint 1740, enabling,for example, positive detection of “foot-down” for use by the controller1762 state machine that controls the apparatus 1700 functions;measurement of applied moment for use by the positive-feedback reflexcontrols employed during powered plantarflexion; and positive detectionof tripping for use by a safety system integrated into the controller1762.

In this embodiment, the structural element 1732 is designed as aflexural element that amplifies the strain fields induced by themedial-lateral moment and axial force applied to the apparatus 1700during operation. The structural element 1732 creates high magnitudestrain fields of opposite sign (differential strain fields) in theregions 1738 and 1742 about the center adaptor mounting hole 1734 when amedial-lateral moment (moment about the x-axis) is applied. Thesedifferential strain fields are not present when only an axial force isapplied. The structural element 1732 includes one strain gage (1782 and1786) bonded to each of the two moment-sensitive regions (1738 and 1742,respectively) on the bottom surface 1733 of the structural element 1732.The gages are applied on opposing sides of a Wheatstone bridge. Thecontroller 1762 is coupled to the Wheatstone bridge to measure thestrains. The strain measurements are used to measure moment on thestructural element 1732. In one embodiment, the sensitivity of themeasurement is approximately in the 0.15 N−m range, where, in thiscontext, sensitivity defines the minimum resolvable change (signal tonoise≈1) when digitally sampled at 500 Hz.

In contrast to the moment induced strains, high strains are introducedby axial forces along the medial-lateral axis in the regions 1746 and1754 around the center adaptor mounting hole 1734. These strains appearin a 0.76 mm thickness region (regions 1746 and 1754) under the slots(1758 and 1770, respectively) machined along the medial-lateral axis.The section above the slot must be thick enough to transfer moment loadwith minimum strain in the thin lower section. The strain magnitude issignificantly diminished in the thin section when a moment-only load isapplied. The structural element 1732 includes one strain gage (1790 and1794) bonded to each of the two axial load-sensitive regions (1746 and1754, respectively) on the bottom surface 1733 of the structural element1732. The gages are applied on opposing sides of a Wheatstone bridge.The controller 1762 is coupled to the Wheatstone bridge to measure thestrains. The strain measurements are used to measure axial force on thestructural element 1732, and consequently, axial force on the lower legmember 1712. The machined slots 1758 and 1770 amplify theaxially-induced strains without compromising the structural integrity ofthe structural element 1732.

Since the structural element 1732 is in the critical chain of structuralsupport between the residual limb socket of the wearer (not shown) andthe apparatus 1700, in one embodiment it is preferably designed towithstand more than 60 N/kg of axial load. In this embodiment, thesensitivity of the axial measurement is in the range of approximately 50N, which is well below the approximately 100 N threshold normally usedin the apparatus 1700 to sense that the apparatus has been placed firmlyon the ground. During calibration of the apparatus 1700 a 2×2sensitivity matrix is determined, enabling true moment and axial forceto be derived from the pairs of strain measurements.

FIG. 17F is an illustration of a cross-sectional view of an alternativemethod for measuring axial force and moment applied to a lower legmember, according to an illustrative embodiment of the invention. Inthis embodiment, the structural element 1732 employs a flexural designthat amplifies displacement of its bottom surface 1733 in such a waythat the axial force and in-plane moment (two-degrees of freedom) can bederived in a redundant fashion. In this embodiment, the apparatus 1700includes a displacement sensing apparatus 1735 for measuring deflectionof the structural element 1732 to determine the moment (torque) andaxial force applied to the lower leg member 1712.

In this embodiment, the displacement sensing apparatus 1735 includes aprinted circuit assembly (PCA) employing one or more displacementsensors 1737 (e.g., contact or non-contact displacement sensors). Thesensors measure, at each sense coordinate, the distance between thesensor 1737 and the bottom surface 1733 of the structural element 1732.

In one embodiment, changes in mutual inductance of coils printed on thePCA with respect to the bottom surface 1733 of the structural element1732 are used to measure the local surface deformation (displacement).In this embodiment, counter-circulating “eddy” currents in thestructural element 1732 serve to reduce the coil inductance inverselywith the distance between the coil and the bottom surface 1733 of thestructural element 1732. Other displacement sensing technologies couldbe employed, including non-contact capacitance and optical sensors orcontact-based sensors that employ force-sensitive resistors, piezo orstrain-gages integral to the PCA. By sampling the array of displacementsensors, the axial force and moments can be estimated using asensitivity matrix that is computed during an off-line calibrationprocess.

In this embodiment, the structural element 1732 is fastened to the lowerleg member 1712 with screws, eliminating the need for the pins 1711employed in the embodiment illustrated in FIG. 17E. The screw fasteningmethod reduces weight and manufacturing complexity. Furthermore, thisfastening method amplifies displacements measured in the center of thestructural element 1732 where the displacement sensing apparatus 1735 islocated. FIG. 17G illustrates how the in-plane moment vector and axialforce may be computed using a circular array of displacement sensors onthe printed circuit assembly. As shown, demodulation of the bias andsinusoidal-like displacement function is used to estimate the moment andforce. Other displacement sensor array configurations could be used as ameans of intrinsic sensing of moment and force.

Moment and force sensing is useful as a means of signaling walking statechanges. In addition, measurement of lower leg member 1712 moment servesas a feedback means by which reflexive behavior is achieved in poweredplantarflexion. When combined with inertial and actuator feedback, theintrinsic moment and force measurements are used to calculate groundreaction force and zero moment pivot, which are useful for tractioncontrol and balance. For these reasons, it is beneficial to package theintrinsic moment and force sensing with the inertial measurement unitand state control processing functions. FIG. 17F shows how thesefunctions could be implemented on a PCA. Such a PCA could be implementedas a sandwich of FR-4 material with a stable core material (Invar forinstance) serving as a stiff interposing substrate between the top-sidedisplacement sensing FR4-based layer and a bottom FR-4-based layer thatincorporates the signal processing layer. Integrating the materials andfunctions in a single assembly eliminates the need for cabling and otherpotentially unreliable means for interconnecting these functions. Suchintegration also allows for a stand-alone tool that can be used byprosthetists to setup a passive prosthetic and study, gait parameters,including energy return and walking statistics

Referring to FIG. 17A, the apparatus 1700 also includes a controller1762 coupled to the linear actuator 1716 for controlling the linearactuator 1716. In this embodiment, the controller is located within ahousing 1764 of the apparatus 1700 to protect it from the environment. Abattery 1768 in the housing 1764 provides power to the apparatus (e.g.,the controller 1762 and various sensors associated with the apparatus1700).

The apparatus 1700 includes an inertial measurement unit 1720 to predictthe inertial pose trajectory of the ankle joint 1740, heel 1772 and toe1776 relative to the previous toe-off position. The inertial measurementunit 1720 is electrically coupled to the controller 1762 and providesinertial measurement signals to the controller 1762 for controlling thelinear actuator 1716 of the apparatus 1700. In one embodiment, theinertial measurement unit 1720 employs a three-axis accelerometer and athree-axis rate gyro. The three-axis accelerometer measures localacceleration along three orthogonal axes. The three-axis rate gyromeasures angular rotation about three orthogonal axes. Through use ofwell-established methods of numerical integration, the position,velocity and pose of any point on the foot structure can be computed.

In some embodiments, the inertial measurement unit 1720 is used todetect the terrain slope and the presence of steps and stairs—therebyenabling optimization of the foot's “angle-of-attack” relative to theunderlying terrain prior to touchdown and the ankle joint's springequilibrium position in the stance phase. In some embodiments, theinertial measurement unit 1720 is used to determine ambulation speed ofthe wearer and conditions of the terrain (features, texture orirregularities of the terrain (e.g., how sticky is the terrain, howslippery is the terrain, is the terrain coarse or smooth, does theterrain have any obstructions, such as rocks)). This enables the wearerto walk confidently on all terrain types. The inertial pose comprisesthe three degree-of-freedom orientation of the lower leg member 1712 ina fixed ground-referenced (world) coordinate frame—often captured as theorientation component of a homogeneous transformation (three unitvectors defining the x, y and z axes in the world reference frame) or asa quaternion; the translation of the ankle joint 1740 in the worldframe; and the velocity of the ankle joint 1740 in the world frame. Inthis embodiment, the inertial measurement unit 1720 is physicallycoupled to the lower leg member 1712. In some embodiments, the inertialmeasurement unit 1720 is coupled to the foot member 1708 of theapparatus 1700.

FIG. 17H is a schematic illustration of a state estimation and actuatorcontroller (state and actuator control PCA-SAC) for use with theapparatus of FIGS. 17A-17G, according to an illustrative embodiment ofthe invention. In this embodiment, the controller 1762 employs dual 40MHz dsPIC (manufactured by Microchip™) processors to control andcoordinate linear actuator 1716 (e.g., rotary motor 504 of FIGS. 5A and5B) and the inertial measurement unit 1720. In this embodiment,space-vector modulation is employed to implement the brushless motorcontrol to create an optimum pulse width modulated drive that maximizesmotor RPM. Space vector modulation is a PWM control algorithm formulti-phase AC generation, in which a reference signal is sampledregularly. PWM of a signal or power source involves the modulation ofthe three-phase motor winding voltage duty cycles (e.g., the rotarymotor 504). After each sampling of the reference signal, non-zero activeswitching vectors adjacent to the reference vector and one or more ofthe zero switching vectors are selected for the appropriate fraction ofthe sampling period in order to synthesize the reference signal.

The controller 1762 receives a variety of input signals, including,inertial pose signals 1781 from the inertial measurement unit 1720,torque and axial force signals 1783 from the structural element 1732strain measurements, ankle joint angle signals 1785 from a hall-effecttransducer located in the ankle joint 1740, motor position signals 1787(quadrature encoder with index and absolute motor position) from theencoder (e.g., encoder 2040 of FIG. 20A), strain signals 1789 from thestrain sensor 1704 (referring to FIG. 18A) of the series elastic member1728, and controller parameters 1791 (e.g., apparatus configurationdata, wearer-specific tuning, firmware updates)). In addition, thecontroller 1762 outputs a variety of signals, including, apparatusperformance data 1793 (e.g., real-time data, error log data, real-timeperformance data), ankle state updates 1795. In addition, the controller1762 outputs commands to the linear actuator 1716 and receives actuatorfeedback signals from the linear actuator 1716 (generally signals 1797),for example, three-phase pulse width modulation signals provided to thepower electronics for the linear actuator 1716, battery power to thelinear actuator 1716, and current feedback measurements and temperaturemeasurements from the linear actuator 1716.

This embodiment uses the sensor feedback to identify state changes asthe apparatus 1700 transitions through the stance-phase and swing-phasestates. By using the redundant and diverse sensor measurements, it alsoidentifies fault conditions and drives the apparatus 1700 into anappropriate safe state. Using an on-board real-time clock, it time-tagsthe faults and stores these into an on-board e²PROM (error log). Thecontents of the error log are retrieved wirelessly by the prosthetistand/or manufacturer service personnel. In this embodiment, the MotorDriver PCA (MD) takes pulse-width modulation (PWM) commands from the SACPCA to switch current into the motor windings. The MD passes sensedcurrent and power information back to the SAC PCA so that it can applyclosed-loop control.

In this embodiment, the IMU PCA is mounted nominally in the Sagittalplane (a local plane parallel to the front of the tibia) and employs athree-axis accelerometer, a dual-axis rate gyro (ω_(z) and ω_(x)) and asingle-axis rate gyro (ω_(y)). In this embodiment, a coordinate framedefinition is used that defines the y-axis as forward, z-axis as up andx-axis defined as the cross-product of the y and z axes (y×z). The IMUreceives state information from the SAC at the system sampling rate of500 Hz. It transmits the ankle state vector—specifically the positionand velocity of the ankle pivot, the position of the heel and theposition of the toe—all with respect to the foot-flat position from theprevious step.

FIGS. 17I and 17J are schematic illustrations of an exemplary electricalequivalent of apparatus 1700 of FIG. 17A. Electrical circuit symbols areused to describe the mechanical elements—a resistor denoting amechanical component with damping torque that is linear with velocity; acapacitor denoting a mechanical component with rotary inertiaproperties; and an inductor denoting a mechanical component with linearspring qualities. With this circuit notation, current corresponds withtorque and voltage corresponds with angular velocity.

The circuit components are defined as follows: J_(shank) is the unknownequivalent inertia of the lower leg member (shank) and residual limbbelow the knee (e.g., inertia of lower leg member 1712 of FIG. 17A);J_(motor) is the equivalent motor and ball-screw transmission assemblyinertia (e.g., inertia of linear actuator 1716 of FIG. 17A); K_(series)^(compression) is the torsional spring constant for the series spring(e.g., passive elastic element 1728 of FIG. 17A) when in compression;K_(series) ^(tension) is the torsional spring constant for the seriesspring when in tension; K₃ is the torsional spring constant for theunidirectional parallel spring (e.g., passive elastic member 1724 ofFIG. 17A); and J_(Ankle) is the rotary inertia of the foot structurebelow the ankle (e.g., foot member 1708 of FIG. 17A). The current(torque) sources within the model are defined as follows: Γ_(Human) isthe unknown torque applied by the wearer's body onto the lower legmember (e.g., lower leg member 1712); τ_(motor) is the torque applied bythe actuator (e.g., linear actuator 1716); and Γ_(shank) is the torquemeasured using the structural element (e.g., structural element 1732 ofFIGS. 17A and 17E).

FIG. 17I illustrates the importance of the series and parallel springsas energy storage elements. Use of the stored energy reduces the powerconsumption that would otherwise be required by the linear actuator. Inaddition, an additional purpose of the K₃ spring is its function as ashunt across the ankle inertia that increases the ankle-springresonance.

FIG. 17J illustrates how sensors have been employed in this embodimentto provide high-fidelity position and force control, and to achieve thesensor redundancy and diversity desirable for delivering an inherentlysafe design. As shown, the ankle joint position, {circumflex over (θ)},is derived from the following:

$\begin{matrix}{\hat{\theta} = {\theta_{motor} - \frac{F_{series}}{K_{series}}}} & {{EQN}.\mspace{14mu} 38} \\{where} & \; \\{K_{series} = \left\{ \begin{matrix}{K_{series}^{compression},} & {{{if}\mspace{14mu}\Gamma_{series}} \geq 0} \\{K_{series}^{tension},} & {{{if}\mspace{14mu}\Gamma_{series}} < 0}\end{matrix} \right.} & {{EQN}.\mspace{14mu} 39}\end{matrix}$

A redundant measure of θ is achieved through use of a Hall-effect angletransducer, thereby providing a verification that the ankle is beingmanipulated properly by the control system. In one embodiment, theHall-effect transducer includes a Hall-effect device located on the SACPCA in the housing 1764 of the apparatus 1700. The transducer alsoincludes a magnet coupled to the foot member 1708. The field effectmagnitude (signal output by the transducer) changes in a known way inresponse to angle joint rotation (i.e., motion of the magnet relative tothe Hall-effect device). The Hall-effect transducer is calibrated duringmanufacturing of the apparatus 1700 by, for example, measuring theoutput of the transducer to known displacements of the Hall-effectdevice relative to the magnet. In other ankle angle measurementembodiments, the mutual inductance measured on a coil on the lower legmember has a known relationship as a function of ankle angle, and theinductance can be calibrated to compute angular displacement in a waythat is not sensitive to the magnetic fields generated by the motor inthe linear actuator or by other stray fields. Also, as shown in FIG.17J, the ankle moment as applied by the wearer is also measured. Thisenables the linear actuator to adapt (e.g., to increase stiffness) toachieve reflex behavior.

FIGS. 18A, 18B, 18C and 18D are illustrations of the passive elasticmember 1728 of FIG. 17A, according to an illustrative embodiment of theinvention. The passive elastic member 1728 provides bidirectionalstiffness and is connected in series with the linear actuator 1716 andthe foot member 1708. The passive elastic member 1728 is coupled at oneend to the second end 1748 of the linear actuator 1716, and at the otherend to the foot member (not shown). The passive elastic member 1728includes a strain sensor 1704 coupled to the passive elastic member 1728for measuring strains in the passive elastic member 1728. In thisembodiment, the strain sensor 1704 is a strain gage whose response iscalibrated to measure the force applied by the linear actuator 1716—andin turn, the moment about the ankle joint 1740 that is applied by thelinear actuator 1716. The strain gage signal is measured using thecontroller 1762 of FIG. 17A.

In this embodiment, the passive elastic member 1724 is a formedcarbon-fiber layup that delivers a desired bidirectional (functions inbending in both directions) normalized drive stiffness. In oneembodiment, the passive elastic member 1724 has a preferred compressionof 14-25 N-m/rad/kg and tension: 4-8 N-m/rad/kg. Biomechanical forcesand torques strongly scale with body mass of a wearer. When scalingprosthetic and orthotic devices, design parameter specifications aretypically normalized. For example, series and parallel elasticity suchdevices can be scaled with body mass, or designed to provide discretevalues that are intended to cover several ranges of body mass. Theranges of compression and tension reflect the variation in torque thatresults from the difference in the linear actuator moment arm to theankle joint across the entire range of rotation-from maximumplantarflexion to maximum dorsiflexion. The series spring constant isoptimized to be relatively non-compliant during swing-phase dorsiflexionposition control (while the spring is in compression) such as when theankle is being repositioned immediately following toe-off in walking.However, some compliance is maintained to isolate the linear actuatorfrom shock loads.

Referring to FIGS. 18C and 18D, high stiffness is achieved in thepassive elastic member 1728 in compression by inserting a dorsiflexionrotation bottom constraint 1708 towards the distal end 1726 of thepassive elastic member (spring) 1728. This restraint reduces theeffective moment arm of the linear actuator 1716 on the bending of theseries spring 1728 during compression (towards dorsiflexion). Intension, the moment arm is effectively increased by placing theplantarflexion top constraint 1716 more towards the proximal end 1730 ofthe spring restraint. With the longer moment arm, the spring beam willbend more freely, thereby reducing the spring constant in tension. Inaddition to the bilateral stiffness characteristics, in someembodiments, the spring constant of the passive elastic member 1728 isoptimized to minimize ball-screw rotational speed By design, thisembodiment of the elastic member 1728 has asymmetricalcharacteristics—delivering higher compliance in tension than incompression. The higher compliance in tension increases the energystorage in the series spring 1728 for use in powered plantarflexion. Theenergy is released in about the first 100 ms involved in poweredplantarflexion, thereby reducing the energy contribution required of thelinear actuator 1716. In embodiments of the invention that use aball-screw transmission assembly in conjunction with a rotary motor forthe linear actuator (e.g., ball-screw transmission assembly 2024 ofFIGS. 20A-20B), this has the added benefit of reducing the requiredoperating speed of the ball-nut assembly portion of the ball-screwtransmission assembly and also the motor drive requirements for therotary motor. The spring catapults the foot member without requiringhigh-speed ball-nut positioning in this case. Optimized values for theseries elasticity are in the range of 3-4 Nm/rad/kg.

FIG. 19A is an illustration of a lower-extremity prosthetic apparatus1900 incorporating a flat series spring 1928, according to anillustrative embodiment of the invention. The apparatus 1900 has amounting interface 1904 making it capable of attaching to acomplementary lower-extremity limb socket member of a wearer. Theapparatus 1900 includes a lower leg member 1912 coupled to the mountinginterface 1904. The lower leg member 1912 is also coupled to a footmember 1908 at an ankle joint 1940 of the apparatus 1900. The anklejoint 1940 permits the foot member 1908 to rotate about the x-axisrelative to the lower leg member 1912. The foot member includes a heel1972 and a toe 1976.

The apparatus 1900 also includes a linear actuator 1916 with a first end1936 and a second end 1948. The first end 1936 of the linear actuator1916 is coupled to the lower leg member 1912. The apparatus 1900 alsoincludes passive elastic member 1928 in series with the linear actuator1916. The passive elastic member 1928 is coupled between the foot member1908 and the second end 1948 of the linear actuator 1916. The passiveelastic member 1928 is coupled to the foot member 1908 at the proximalend 1930 of the passive elastic member 1928. The distal end 1926 of thepassive elastic member 1928 is coupled to the second end 1948 of thelinear actuator 1916. The linear actuator 1916 applies torque about theankle joint 1940.

The apparatus 1900 also includes a controller 1960 coupled to the linearactuator 1916 for controlling the linear actuator 1916. In thisembodiment, the controller 1960 is located within a housing 1964 of theapparatus 1900 to protect it from the environment; however, a portion ofthe housing is removed in this figure to expose the contents within thehousing). A battery 1968 coupled to the apparatus 1900 provides power tothe apparatus 1900 (e.g., the controller 1960 and various sensorsassociated with the apparatus 1900).

The passive elastic member 1928 of FIG. 19A is a flat spring (e.g.,fabricated with water-cut equipment). A flat spring reduces the cost ofthe passive elastic member 1900 and makes it easier to configure thespring constant to align with the body mass of the wearer. In oneembodiment, the spring is split longitudinally (along the y-axis) toreduce the out-of-plane moment on the components of a ball-nut (see,e.g., FIGS. 20A and 20B) of the linear actuator 1916 due to lack ofparallelism between the rotation axis of the ball-nut and the seriespassive elastic member 1928. In this embodiment, no strain sensing isemployed in the actuator torque feedback loop. Rather, the torquetransmitted through the spring is estimated by multiplying the knownspring constant of the flat spring by the measured spring deflection(difference between measured ankle joint 1940 angle, θ and the angle, β,kinematically defined as the ankle joint 1940 angle that would resultfrom a specific ball-nut position along the screw when the springdeflection is zero.

FIGS. 19B and 19C are illustrations of an alternative two-pieceseries-elastic spring of a prosthesis apparatus 1900, according to anillustrative embodiment of the invention. The apparatus 1900 has amounting interface 1904 making it capable of attaching to acomplementary lower-extremity limb socket member of a wearer. Theapparatus 1900 includes a lower leg member 1912 coupled to the mountinginterface 1904. The lower leg member 1912 is also coupled to a footmember 1908 at an ankle joint 1940 of the apparatus 1900. The anklejoint 1940 permits the foot member 1908 to rotate about the x-axisrelative to the lower leg member 1912. The foot member includes a heel1972 and a toe 1976. The apparatus 1900 also includes a linear actuator1916 with a first end (not shown) and a second end 1948. The first endof the linear actuator 1916 is coupled to the lower leg member 1912. Theapparatus 1900 also includes a coupling member 1988 (e.g., bracket) thatcouples the foot member 1908 to the lower leg member 1912 at the anklejoint 1940 with a bearing that allows the foot member 1908 to rotateabout the x-axis of the ankle joint 1940.

The apparatus 1900 also includes passive elastic member 1928 in serieswith the linear actuator 1916. Referring to FIG. 19C, the passiveelastic member 1928 has two member sections (e.g., beam-like sections)1994 and 1996. The elastic member 1928 also has a first end 1962 on thefirst member 1994 and a second end 1980 on the second member 1996. Theelastic member 1928 also has an intermediate location 1996 at which thetwo members 1994 and 1996 meet and at which the two members 1994 and1996 pivot with respect to each other around the x-axis. As the secondmember 1996 pivots towards the first member 1994, the elastic memberstores energy in compression during dorsiflexion (shown by the arrow1992).

The first end 1962 of the elastic element 1928 is coupled to the secondend 1948 of the linear actuator 1916 with a bearing that allows forrotation about the x-axis. The second end 1980 of the elastic element1928 couples to a location on the coupling member 1988 with a bearingthat allows for rotation about the x-axis.

Exemplary Linear Actuator

FIGS. 20A and 20B are illustrations of a linear actuator 2000 for use invarious lower-extremity prosthetic, orthotic, and exoskeleton apparatus,according to an illustrative embodiment of the invention. FIG. 20A is aperspective view of the linear actuator 2000. FIG. 20B is across-sectional view of the linear actuator 2000. The linear actuator2000 can be used as, for example, the linear actuator 1716 of apparatus1700 of FIG. 17A or apparatus 400 of FIG. 4. The actuator 2000 includesa motor 2004 and screw transmission assembly 2024 (in this embodiment,it is a ball-screw transmission assembly, also referred to as aball-screw assembly) for delivering linear power along the A axis. Thescrew transmission assembly 2024 functions as a motor drive transmissionto translate rotational motion of the motor 2004 to linear motion. Inone embodiment, the ball-screw transmission assembly 2024 is a customball-screw transmission assembly manufactured by Nook Industries(offices in Cleveland, Ohio). The custom ball-screw transmissionassembly has the following specifications: 14 mm×3 mm pitch screw, 4000N of thrust at 150 mm/s, and an L1 rated life in the instant applicationof 5 million cycles. In some embodiments, the screw transmissionassembly is a lead-screw transmission assembly (also referred to as alead-screw assembly).

The actuator 2000 includes a rotary motor 2004 that has a motor shaftoutput 2008. The motor shaft output 2008 has a pulley 2032 coupled(e.g., welded) to the motor shaft output 2008. In one embodiment, therotary motor 2004 is a high-speed brushless motor (model EC30 motormanufactured by Maxon Motor AG, Maxon Precision Motors, Inc. withoffices in Fall River, Mass.). The motor 2004 includes an inductiveincremental-absolute angular encoder 2040 that is integrated into themotor 2004 to for determining angular alignment between the rotor andstator of the rotary motor 2004. The encoder 2040 also provides aposition feedback signal necessary to control the screw 2060 position ofthe linear actuator 2000 and to provide for “instant-on” motorcommutation and redundant position feedback monitoring.

The inductively-coupled encoding elements of the encoder 2040 enable thesystem to determine the absolute rotor-stator alignment (with, forexample, 10 bits of resolution per revolution) simultaneously withhigh-precision incremental rotor position feedback. By cross-checkingthese redundant feedback elements it is possible to minimize thepossibility that an encoder malfunction can cause ankle controlinstability. The incremental encoder achieves less than 300 grad ofrun-out so as to eliminate the sensed velocity fluctuations when theball-screw transmission assembly 2024 (see below) is operating atconstant-speed. As a result, less torque variation is applied by theactuator 2000.

The rotary motor 2004 also includes an integral motor heat-sink 2048. Inone embodiment, the heat-sink 2048 draws heat out of the windings of themotor 2004, enabling a wearer to walk at peak levels of non-conservativework without exceeding motor coil temperature limits (typically 160°C.). Motor heating arises due to resistive losses (i²R losses) in themotor 2004 as the linear actuator 2000 delivers thrust force. As thecoil temperature rises, the coil resistance rises at the rate of 0.39%/°C., thereby further increasing the coil temperature. In addition, themotor K_(t) (a measure of torque as it scales with motor current)typically drops by nearly 20% as the coil temperature increases to itslimit. This requires additional current consumption to perform the sameamount of work, further driving up the coil temperature. The heat-sinkin the linear actuator 2000 reduces coil temperature rise by over 40%.Because the wear out phenomenon that drives premature failure of motorwinding insulation and motor bearing reduces in effect by a factor of 2×for every coil temperature reduction of 10° C., the motor life increasessignificantly if lower motor coil operating temperatures are maintained.And, using this intrinsic coil temperature sensing method, the motor canbe protected from exceeding the absolute maximum rating of 160° C. bysimply reducing powered plantarflexion power (currents) as the maximumrating is approached, and ultimately, shutting off battery power when apredefined limit of, for example, 150° C. is reached.

Robotic prostheses typically employ compact light-weight motor drives todeliver power in bursts to the affected limb. In some scenarios, thepower bursts may be applied repetitively and at high rate over extendedperiods of time. The motor copper and eddy current losses will cause anexcessive accumulated heating effect that causes the motor windingtemperature to rise. Since the copper winding resistance increases withtemperature (0.39%/° C.), the copper losses will increase therebyamplifying the heating effect. A critical winding temperature limit cansometimes be reached in which further temperature rise will causepermanent damage to the motor. Sensing when this temperature limit isreached is preferably accomplished by the control system.

Two conventional methods may be used to prevent or detect when thecopper winding temperature limit is or will be reached. In the first,the copper and eddy current losses are computed while the control systemoperates. These are used to drive a thermal model of the windings sothat the winding temperature can be estimated. Sometimes the ambienttemperature is measured in order to yield a better winding temperaturemeasurement. An advantage of this method is that it is cheap toimplement. The disadvantage is that the coil temperature model is hardto obtain and to calibrate. Also, it is often difficult to make a goodmeasurement of the ambient temperature around the motor, causing thewinding temperature measurement to be in error.

In the second method, sometimes combined with the first, the casetemperature of the motor is measured with a thermistor applied to theoutside of the case, or inside the motor. The advantage of this is thatit provides a direct measurement. The disadvantage is that it onlymeasures at one point and the application of the sensor is expensive andoften unreliable.

A more preferred approach is to both detect the temperature and tomitigate the potential overheating condition. In this, we measure themotor winding resistance on every step at a point during the walkingcycle when we can briefly hold the ankle at a fixed position (this toeliminate back-emf effects on the resistance calculation) to make themeasurement. In one embodiment, coil temperature is determined byapplying a fixed current (alternatively fixed voltage) to the motorwinding and measuring the corresponding voltage (alternatively current)in the winding. To increase the accuracy, we apply the voltage (orcurrent) in both the forward and reverse direction and measure thedifference in current (or voltage).

Since the motor drive electronics employs PWM current control methods,all the infrastructure to make this measurement exist. By noting thepercentage difference between this winding resistance and that when theankle is at rest (a calibration constant) we can estimate the windingresistance in-situ without cost. In a typical servo system thismeasurement cannot be made because the actuator must be continually inclosed-loop control. But in the ankle prosthesis, there are times (swingphase) when the ankle position does not need to sustain the precisioncontrol over the 5 milliseconds typically required to make themeasurement. Once the winding temperature is calculated in this way, thecontrol system can detect when the windings are approaching the criticaltemperature. During these times, the drive power available for walkingis reduced or eliminated altogether until the temperature is reduced toa safe level.

In some embodiments, the output of the temperature sensor 2052 isprovided to a controller (e.g., the controller 1762 of FIG. 17A) tocontrol torque output by the linear actuator 2000 based on thetemperature of the motor 2004.

A belt 2012 couples the pulley 2032 to the threaded shaft 2060 of aball-screw transmission assembly 2024 such that rotational motion of themotor shaft output 2008 is translated to a linear motion of the ball-nutassembly 2036 portion of the ball-screw transmission assembly 2024. Insome embodiments, two or more belts are applied in parallel, each withan ability to drive the linear actuator 2000 ball-screw transmissionassembly 2024 by itself, so as to enable the linear actuator 2000 tosurvive a single belt breakage failure. In such an event, belt breaksensor 2056 senses the condition and validates belt integrity duringoperation (e.g., during each gait cycle of a wearer using a prosthesis).

In one embodiment, an optical sensor (e.g., a thru-beam sensor) is usedas the belt break sensor and an output signal of the optical sensorchanges in a known manner when a belt breaks. In another embodiment ofthe invention, a capacitive sensor is used as the belt break sensor andan output of the capacitive sensor changes in a known manner when a beltbreaks.

In one embodiment, the pulley 2032 and belt(s) are not used as theapparatus for converting rotary motion to a linear motion. Rather, a setof traction wheels are used as the transmission apparatus. In thisembodiment, the threat of belt failure is thereby eliminated.

In one embodiment, in the event of a belt break, a controller of theapparatus in which the linear actuator 2000 is used (e.g., controller1762 of apparatus 1700 of FIG. 17A), changes the position of the footmember relative to the lower leg member to a safe position that enablesthe apparatus to operate as a passive ankle prosthesis until the linearactuator 2000 is repaired. In one embodiment, the controller shortsthree electrical leads of the rotary motor 2004 in response to the beltbreakage sensor detecting a failure of one or more of the plurality ofbelts. Shorting the three-phase electrical input leads to the motor 2004introduce a viscous drag on the motor shaft output 2008. While walking,the viscous drag holds roughly fixed the rotor shaft output so that theapparatus operates as a passive prosthesis. However, the apparatus canbe moved slowly in a way that enables it to move to a non-fixedequilibrium position when standing or sitting. Each input lead isshorted to ground by its own individual switch.

In one embodiment, the switches are operated by a rechargeable battery(a separate battery from the primary battery used to operate theapparatus). By using a separate battery, the switches would short theinput leads (and place the apparatus into a safe mode) even if a failureoccurred (or the primary battery was depleted).

In one embodiment, the threaded shaft 2060 includes a hollowed outportion that contains a noise damping material to reduce the noisegenerated by the actuator 2000 and the apparatus within which theactuator 2000 is used. In one embodiment, the threaded shaft 2060 is 14mm diameter stainless steel shaft 8.7 mm diameter bore that extends 64mm of the length of the shaft, filled with ISODAMP® C-1002 acousticdamping material manufactured by 3M (with offices in St. Paul, Minn.).

The actuator 2000 also includes a radial and thrust bearing 2028 thatsupport the belt 2024 tension due to the rotary motor 2004 and thethrust force of the screw 2036. Loads due to the belt tension and thrustforce are present both statically and during the gait cycle.

The ball-nut assembly 2036 includes one or more recirculatingball-tracks 2042 that retain a plurality of ball bearings; thecombination of which support the linear motion of the ball-nut assembly2036. In one embodiment, five ball-tracks are used. The actuator 2000includes a coupling element 2020 (e.g., the second end 1748 of thelinear actuator 1716 of FIG. 17A) that couples the actuator 2000 to, forexample, a passive elastic member of a foot member of a prostheticapparatus (e.g., passive elastic member 1724 of FIG. 17A).

FIG. 21 is an illustration of a perspective view of a linear actuator2100 for use in various lower-extremity prosthetic, orthotic, andexoskeleton apparatus, according to an illustrative embodiment of theinvention. The linear actuator 2100 can be used as, for example, thelinear actuator 1016 of apparatus 1000 of FIG. 17A or apparatus 400 ofFIG. 4. The linear actuator 2100 is a variation of the actuator 2000 ofFIGS. 20A and 20B.

The actuator 2100 includes a rotary motor 2004 that has a motor shaftoutput 2008. The motor shaft output 2008 has a pulley 2032 welded to themotor shaft output 2008. The motor 2004 includes an inductiveincremental-absolute angular encoder 2040 that is integrated into themotor 2004 to for determining angular alignment between the rotary motor2004 rotor and stator. The rotary motor 2004 also includes an integralmotor heat-sink 2048.

Two belts 2104 a and 2104 b are used in parallel, rather than the singlebelt 2012 of FIGS. 20A and 20B. Each belt has the ability to drive thelinear actuator transmission by itself with 1.5× margin on beltbreakage, so as to enable the linear actuator 2100 to survive a singlebelt breakage failure. In one embodiment, in the event of a belt break,a controller of the apparatus in which the linear actuator 500 is used(e.g., controller 1762 of apparatus 1700 of FIG. 17A) moves the ankle toa safe position in a way that would enable the apparatus to operate as apassive ankle prosthetic until the linear actuator 500 is repaired. Inone embodiment, the controller shorts three electrical leads of therotary motor 504 in response to the belt breakage sensor detecting afailure of one or more of the plurality of belts. In such an event, oneor more belt break sensors would sense the condition and move the ankleto a safe position in a way that would enable the system to operate as apassive ankle prosthesis until the linear actuator is repaired.

The two belts 2104 a and 2104 b couple the pulley 532 to a threadedshaft of a ball-screw transmission assembly (e.g., threaded shaft 2060of FIG. 20B) such that rotational motion of the motor shaft output 2008is translated to a linear motion of the ball-nut assembly 2036 portionof the ball-screw transmission assembly. The actuator 2100 also includesa radial and thrust bearing 2028 that support the tension in belts 2104a and 2104 b due to the rotary motor 2004 and the thrust force of thethreaded screw. Loads due to the belt tension and thrust force arepresent both statically and during the gait cycle.

The ball-nut assembly 2036 includes recirculating ball-tracks thatretain a plurality of ball bearings; the combination of which supportthe linear motion of the ball-nut assembly 2036. The actuator 2100includes a coupling element 2020 (e.g., the second end 1748 of thelinear actuator 1716 of FIG. 17A) that couples the actuator 2100 to, forexample, a passive elastic member of a foot member of a prostheticapparatus (e.g., passive elastic member 1724 of FIG. 17A).

The actuator 2100 also includes a ball-screw assembly seal 2108. Theball-screw assembly seal 2108 protects the screw from, for example,contaminants (e.g., sand, dirt, corrosive materials, sticky materials).Such contamination would cause the design life of the actuator to becomeindeterminate.

Exemplary Lower-Extremity Orthotic (Wearable Robotic Knee Brace)

FIGS. 22A, 22B and 22C are schematic illustrations of a lower-extremityorthotic or exoskeleton apparatus 2200 (wearable robotic knee brace),according to an illustrative embodiment of the invention. The apparatus2200 is a knee-brace that augments the wearer's lower-extremityfunction. FIG. 22A is a top view of the apparatus 2200. FIG. 22B is aside view of the apparatus 2200. FIG. 22C is a view of the interiorportion of a knee joint drive assembly 2204 of the apparatus 2200.Typical use cases for the apparatus 2200 include, for example, metabolicaugmentation, permanent assistance for wearers with a permanent limbpathology, or rehabilitation for wearers with temporary limb pathology.

An example of a metabolic augmentation use case involves, for example,wearers (e.g., soldiers or other personnel) that need to traverse heavyterrain for extended periods at high speed while carrying heavy loads.In this use case, the knee brace apparatus 2200 augments the wearer'sown abilities. An example of a permanent assistance use involves awearer that suffers from a permanent limb pathology (e.g., knee tendonor meniscus degeneration) with no possibility for rehabilitation. Inthis use case, the knee brace apparatus 2200 provides permanentassistance to the wearer. An example of a use case involvingrehabilitation for wearers with temporary limb pathology involves awearer recovering from injury or other temporary condition. In this usecase, the knee brace apparatus 2200 functions as a programmabletelerobotic tool deployed by a physical therapist to acceleraterecovery—through progression of kinesthetic rehabilitation and graduallydecreasing assistance while the muscle memory and strength recover. Inanother embodiment, the method includes specifying a physical therapyprotocol defining a level of assistance performed by the apparatus onthe wearer over a period of time and reducing the level of assistanceperformed by the apparatus on the wearer to assist in rehabilitation ofthe limb pathology. In some embodiments, the level of assistanceperformed by the apparatus is reduced based on impedance and torquecontribution of the wearer to the apparatus.

Referring to FIGS. 22A and 22B, the apparatus 2200 includes a lower legmember 2216 (also referred to as a drive arm), a thigh member 2228, alower leg cuff 2208 and an upper leg cuff 2212. The lower leg cuff 2208is coupled to the lower leg member 2228. The lower leg cuff 2208attaches the apparatus 2200 to the lower leg of the wearer. The upperleg cuff 2212 is coupled to the thigh member 2228. The upper leg cuff2212 attaches the apparatus 2200 to the thigh of the wearer. Theapparatus 2200 includes a knee joint 2232 for connecting the thighmember 2228 to the lower leg member 2216. The knee joint 2232 (e.g., arotary bearing) permits the lower leg member 2216 to rotate about thex-axis relative to the thigh member 2228.

Referring to FIG. 22C, the knee joint drive assembly 2204 includes alinear actuator that drives the knee joint drum 2232 through a beltdrive transmission 2236. The linear actuator is a rotary motor 2240(e.g., brushless motor) and ball-screw transmission assembly 2244 (e.g.,the motor 2004 and ball-screw transmission assembly 2024 of FIGS. 20Aand 20B). In the apparatus 2200, rotational motion of the motor shaftoutput 2256 of the motor 2240 is translated to a linear motion of theball-nut assembly 2248 portion of the ball-screw transmission assembly2244. The motor shaft output 2256 has a pulley 2260 coupled (e.g.,welded) to the motor shaft output 2256. The motor 2240 includes aninductive incremental-absolute angular encoder 2264 that is integratedinto the motor 2240 for determining angular alignment between the rotorand stator of the rotary motor 2240. The encoder also provides aposition feedback signal necessary to control the screw 2252 position ofthe ball-screw transmission assembly 2244 and to provide for“instant-on” motor commutation and redundant position feedbackmonitoring.

A belt 2268 couples the pulley 2260 to the threaded shaft 2252 of theball-screw transmission assembly 2244 such that rotational motion of themotor shaft output 2256 is translated to a linear motion of the ball-nutassembly 2248 portion of the ball-screw transmission assembly 2244.

In one embodiment, the threaded shaft 2252 includes a hollowed outportion that contains a noise damping material to reduce the noisegenerated by the knee joint drive assembly 2204. The knee joint driveassembly 2204 also includes a radial and thrust bearing 2272 thatsupport the belt 2268 tension due to the rotary motor 2240 and thethrust force of the screw 2252. Loads due to the belt tension and thrustforce are present both statically and during the gait cycle.

The knee joint drive assembly 2204 also includes a spring 2280 forseries elasticity, spring cage 2284, drive belt 2236 and a springcage/belt connection 2288. In some embodiments, a drive band (e.g., thinpiece of spring steel) is used in place of the drive belt 2236. In someembodiments, a drive cable (e.g., loop of stranded material) is usedinstead of the drive belt 2236. Spring 2280 is a series passive elasticelement, functioning in the same manner as the series elastic springelement 1728 of FIG. 17A. The spring cage 2284 provides a closed volumein which the spring 2280 is located. The ball-nut transmission assembly2248 is coupled to the screw 2252. The ball-nut assembly 2248 is alsocoupled to the drive belt 2236. Linear motion of the screw 2252 causes alinear motion in the ball-nut assembly 2248. The linear motion in theball-nut assembly 2248 causes a linear motion in the drive belt 2236.The linear motion of the drive belt 2236 drives the knee joint 2232.

The apparatus 2200 includes a controller 2292 (e.g., a printed circuitassembly that incorporates the linear actuator 2204, state and inertialmeasurement unit 2294 (e.g., inertial measurement unit 1720 of FIG. 17A)control and processing functions) to drive and control the operation ofthe apparatus 2200. Referring to FIG. 22B, the apparatus 2200 alsoincludes a torque sensor 2220 coupled to the lower leg member 2216 tomeasure the torque applied to the lower leg member 2216 by the kneejoint drive assembly 2204. The sensor 2220 serves as the feedbackelement in the control loop of the controller 2292 to achieve highfidelity closed loop position, impedance and torque (for reflex) controlof the knee joint 2232. In one embodiment, an array of force-sensitivetransducers are embedded within the cuff structure to provide forcemeasurements used to achieve rapid, biomimetic response.

In some embodiments, the motor angle sensor (e.g., encoder 2264)measures motor position and the controller controls the rotary motor tomodulate position, impedance and torque of the knee joint 2232 based onthe motor position.

In some embodiments, the apparatus 2200 includes an angle sensor fordetermining position of the drum 2232 of the belt drive transmissionrelative to the output of the motor drive transmission and thecontroller controls the rotary motor for modulating impedance, positionor torque based on the position. In some embodiments, the apparatus 2200includes a displacement sensor for measuring displacement of a seriesspring in the motor drive transmission for determining force on theseries spring and the controller controls the rotary motor formodulating impedance, position or torque based on the force on thespring. In some embodiments, the inertial measurement unit 2294 iscoupled to the thigh member or lower leg member for determining aninertial pose of the lower leg member and the controller controls therotary motor for modulating impedance, position or torque based on theinertial pose. In some embodiments, the torque sensor 2220 measures thetorque applied to the lower leg member by the belt drive transmissionand the controller controls the rotary motor for modulating impedance,position or torque based on the torque applied to the lower leg member.In some embodiments, the apparatus 2200 includes an angle sensor fordetermining an angle between the thigh member and lower leg member andwherein the controller controls the rotary motor for modulatingimpedance, position or torque based on the angle between the thighmember and lower leg member.

In some embodiments, the apparatus 2200, instead of a motor drivetransmission, the apparatus includes a screw transmission assemblycoupled to the motor shaft output for converting the rotary motion ofthe motor shaft output to a linear motion output by the screwtransmission assembly. In addition, the drive transmission assemblycoupled to the output of the motor drive transmission is a redundantbelt, band or cable drive transmission coupled to the screw transmissionassembly to convert a linear motion output by the screw transmissionassembly to a rotary motion for applying torque to the knee joint torotate the lower leg member with respect to the thigh member.

Unlike the prosthetic apparatus 2000 of FIG. 20A, the knee braceapparatus 2200 operates in parallel with human actuation. In metabolicaugmentation and replacement applications, the knee brace control systemwill supply all of the impedance and torque needs within the gait cycle.It is desirable for the wearer to be able to walk all day withoutgetting tired and without exertion on the augmented side(s) of the body.In rehabilitation applications, the knee-brace apparatus 2200 suppliesonly a programmed percentage of the impedance and torque. In suchapplications, the knee-brace apparatus 2200 serves as a teleroboticextension of the physical therapist supervising the wearer'srehabilitation.

In one embodiment of the knee brace control system, the physicaltherapist creates a protocol to be executed telerobotically by the kneebrace over a period of time between therapist visits. Using a wirelessinterface, patient performance can be fed back to the physicaltherapist, thereby achieving telepresence. The protocol specifies therate at which the assistance diminishes over time. As the knee braceapparatus reduces assistance, the knee brace apparatus would compute viabiomechanical models the impedance and torque contribution by thewearer—reducing assistance in accordance with the improved response tomaintain the desired net biomimetic response. The biomechanical modelswould involve solving the inverse dynamics of the knee—incorporatinginertial rotation and acceleration of the lower leg member, thigh memberand torso. This six degree-of-freedom information would be derived fromthe inertial measurement unit in the thigh member and knee joint angulardisplacement. The zero-velocity update for the inertial measurement unitwould be accomplished similarly as described herein.

For research/diagnostic purposes, a Wi-Fi (802.11 compatible) module maybe employed to communicate state variables within the control system toan external PC. For use by a prosthetist, a Bluetooth module may beemployed to enable a PDA or phone to monitor, tune and configure theAnkle for the specific needs of an amputee.

It is preferable for the Ankle-Foot System to achieve biomimeticbehavior on all terrain. As shown in FIG. 28, biomechanics studies haveshown that in level-ground walking the normalized ankle torquecorrelates strongly with ankle joint angle and that the normalized peakankle torque is maintained at all walking speeds. This embodiment of theRobotic Ankle-Foot System was designed to be scalable—to employ a familyof interchangeable spring and K₃ ramp kits that can be tailored to sizeand weight of the amputee. Further, through creative use of theactuator, the Γ-θ curve can be further tailored to the amputee—adjustingthe net work applied by the 1-2-3 chain and the 3-8-1 chain as afunction of self-selected walking speed and terrain.

The Robotic Ankle-Foot System can be configured to achieve a Γ-θcharacteristic similar to that shown in the biomechanics study. Threeparameters, K_(P)—the stiffness applied during plantar flexion(heel-strike to foot-flat), K_(D)—the stiffness applied duringcontrolled dorsiflextion, and the degree of exponential hardening(power) can be tailored to create true biomimetic behavior onlevel-ground if desired.

The parameters that adjust the Ankle-Foot System performance andbehavior as described above are preferably adjustable via the wirelessinterface.

Balance Using Ground Reaction Forces and Zero Moment Pivot

FIG. 23A illustrates the generic problem of achieving balance on anincline of variable (positive or negative) slope. The problem appears toinvolve a multi-link, “inverted pendulum” problem, amenable to anon-linear feedback control implementation. In such solutions, knowledgeof the link angles and the mass properties of the links (in this case,leg segments, torso, head and arms) are used to explicitly stabilize themulti-link system. But such explicit inputs are not contained withinmost embodiments of a lower-extremity prosthetic, orthotic orexoskeleton apparatus and would therefore be difficult if not impossibleto implement and package reliably on the wearer. Further, in someinstances, the wearer will have one intact leg, so part of thestabilization will be achieved outside of the lower-extremityprosthetic, orthotic or exoskeleton apparatus, wherein thelower-extremity prosthetic, orthotic or exoskeleton apparatus augmentsthe function of the intact leg.

In addition, FIG. 23B shows that there is a continuum of acceptablesolutions to the balance problem. Specifically, there are an infinitenumber of bent-knee solutions that are entirely acceptable and evendesirable depending on human intent (e.g., picking up heavy luggage orboxes or to achieve balance while playing a game). So we see that thedesired solution will employ intrinsic (to the lower-extremityprosthetic, orthotic or exoskeleton apparatus) sensing that complementsthe intact balance-producing body components to achieve equilibrium inalignment with human intent.

The solution employed in some embodiments of the lower-extremityprosthetic, orthotic or exoskeleton apparatus uses a simplifiedrepresentation of the problem as modeled in FIG. 23C. In thisrepresentation, intrinsic sensing of lower leg member inertial state,ankle joint angle and inertially-referenced, ground reaction force, areused as the stabilization feedback that drives ankle torque (e.g.,torque provided to the ankle joint by a linear actuator of a prostheticapparatus). The body is modeled as a series of masses (only one shown inthe figure) on a massless, thin, buckling beam with time-variablestiffness and moment-of-inertia.

Balance is achieved based on the following details. A desirableequilibrium is achieved when the following conditions are satisfied:

-   -   1. ^(W)F_(GRF) aligns with World z;    -   2. The line connecting the zero moment pivot and the ankle joint        aligns with the World z unit vector; and    -   3. All time derivatives of the inertial lower leg member angle,        γ, and ankle joint angle, θ, are zero.

A feedback control law is then derived that drives each of theseconditions into equilibrium based on the following:τ_(ankle) ≈−Î(t)_(CG) k *(s)[^(ω)γ_(CoP) ^(yω) F _(GRF) ^(y)γ]^(T)  EQN.40wherek *(s)=[k _(r)(s)k _(F)*(s)k _(γ)*(s)]  EQN. 41optimizes the quadratic cost index, J, where

$\begin{matrix}{J = {\int_{0}^{\infty}{\left( {\tau_{ankle}^{2} + {{{{\underset{\_}{k}}^{T}\left\lbrack {\gamma\overset{.}{\gamma}\overset{¨}{\gamma}} \right\rbrack}^{T}\left\lbrack {\gamma\overset{.}{\gamma}\overset{¨}{\gamma}} \right\rbrack}^{T}\underset{\_}{k}}} \right)\ {\mathbb{d}t}}}} & {{EQN}.\mspace{14mu} 42} \\{and} & \; \\{\underset{\_}{k} = \left\lfloor \begin{matrix}k_{\gamma} & k_{\overset{.}{\gamma}} & k_{\overset{¨}{\gamma}}\end{matrix} \right\rfloor} & {{EQN}.\mspace{14mu} 43}\end{matrix}$where the components of k are chosen to emphasize link angle dynamiccontributions to the cost index. In this embodiment, the control lawsolution is provided by the linear-quadratic regulator (LQR)methodology. In layman's terms this means that the settings of a(regulating) controller governing either a machine or process are foundby using the above mathematical algorithms and minimizing a costfunction with weighting factors supplied by a human. The “cost”(function) is often defined as a sum of the deviations of keymeasurements from their desired values. In effect this algorithmtherefore finds those controller settings that minimize the undesireddeviations, for example, deviations from desired work performed by aprosthesis for the wearer. Often the magnitude of the control actionitself is included in this sum as to keep the energy expended by thecontrol action itself limited. In effect, the LQR algorithm optimizesthe controller based on an engineer's specificaiton of the weightingfactors. The LQR algorithm is, at its core, just an automated way offinding an appropriate state-feedback controller.

Use of the quadratic cost index is not required; however, in oneembodiment, use of the quadratic cost index as an optimization criterioncreates an objective framework for analysis and for in-officecustomization for wearers of the lower-extremity prosthesis to achievean acceptable feel as the system works to maintain the wearer'sequilibrium on different terrain. It is not uncommon to find thatcontrol engineers prefer alternative conventional methods like fullstate feedback (also known as pole placement) to find a controller overthe use of the LQR algorithm. With these the engineer has a much clearerlinkage between adjusted parameters and the resulting changes incontroller behaviour.

Wearer Assist in Getting Up from a Chair

FIGS. 24A, 24B and 24C illustrate a method for applying a balancingcontrol law to assist a wearer of a lower-extremity prosthetic apparatusin getting up from a chair, according an illustrative embodiment of theinvention. The Timed Get Up and Go (TUG) is often used as anexperimental means to evaluate dynamic and functional balance. Wearersare given a verbal instruction to stand up from a chair, walk 3 meters,cross a line marked on the floor, turn around, walk back, and sit down.To achieve good “TUG” performance, leg prostheses often have a “standup” and “sit down” button to create the behavioral context for theprosthesis' control system. In the lower-extremity prosthetic apparatusincorporating principles of the present invention, in one embodimentthere is no explicit requirement to set behavioral context by, forexample, pushing a button. Sitting, standing up and sitting downbehavioral context is identified by the intrinsic sensors of theprosthetic apparatus. Control behavior during standing and sitting issimply part of maintaining the wearer's equilibrium.

FIGS. 24A, 24B and 24C illustrate how the intrinsic balance controlalgorithm works to augment the wearer as she stands up from a chair.Referring to FIG. 24A, initiation of the sitting to standing transitioninvolves three states. In the first, the foot is off the ground or onlylightly touching it. The prosthetic apparatus (e.g., apparatus 1700 ofFIGS. 17A-17E) knows the mass of the wearer; the inertial orientation ofthe lower leg member and foot member; and the ground reaction force (asdetermined, for example, with respect to FIG. 11A). The apparatustherefore “knows” or senses that the wearer is sitting. As the wearerbegins to stand up, the increase in ground reaction force is noted andthe state of the foot (foot flat) is known via the inertial measurementunit measurements and ankle joint angle sensor measurements. Theintrinsic balance control law execution begins. During this secondstate, the disequilibrium sensed by the imbalance in the ground reactionforce is used to drive the lower leg member (e.g., driven forward by thecontroller 1762 commanding the linear actuator 1716 to increase thetorque applied to the ankle joint 1740) forward as a means of pullingthe torso (center-of-mass) over the ankle joint.

Referring to FIG. 24B, the intrinsic balance control continues to drivethe wearer into equilibrium in front of the chair. FIG. 24C shows thewearer in mid-stance equilibrium, ready to begin walking if desired. Asshown, wearer intent, and more specifically the sitting/standingbehavioral context can be derived by sensing that is intrinsic to theprosthetic apparatus. The implementation cost and complexity of explicitcontext switching (pressing of buttons) is thereby avoided. Theprosthetic apparatus complements and augments the body function in anatural way.

The ankle torque induced by the ground reaction force (GRF) is apreferred way to achieve exponential hardening during mid-stance. Unlikeuse of the torque on the lower leg (e.g., torque measured using thestructural element 1732 of FIG. 17A), the GRF-computed ankle torquemeasures the torque applied by the ground on the ankle joint. The GRF isoften measured by force plates in gait research settings and is therebyused as a measure of how an intact ankle interacts with the ground whilewalking. The GRF establishes what is the biomimetic ankle behavior indifferent terrain contexts. A benefit of using the GRF as the means bywhich to achieve exponential hardening is the ease by which performancecan be measured relative to biomimetic references. Further, use of thismeasure ensures that invariance to terrain orientation since it derivesfrom intrinsic inertial sensing (e.g., using the inertial measurementunit 1720 of FIG. 17A).

Optimization Methods

FIGS. 25A and 25B are schematic illustrations for controlling alower-extremity apparatus based on a stochastic optimization of 1) thetransition work, W_(t), performed to transfer weight from the trailingleg to the leading leg during the double-support phase of the gait cycle2) minimizing hip impact force and force rate or 3) minimizing acombination of both cost (objective) functions FIG. 25A illustrates thesimplified model used to calculate transition work. FIG. 25B illustratesthe simplified model used to calculate hip impact force and force rate.

The term stochastic denotes that the optimization minimizes the expectedvalue of the objective function subject to hip impact force and forcerate constraints, assuming probability (likelihood) functions for humanintent; biomechanical feedback (including walking speed); terraincontext, and terrain property. The optimization is achieved throughmodification of impedance, torque and position control parameters withinthe control algorithms. Practically speaking, the transfer energy isminimized, and the hip impact force constraints satisfied, by minimizingthe negative impact of foot strike forces and maximizing the positiveimpact of reflex-induced ground forces on the hybrid system energy.

The optimization described above can be implemented in real-time byintroducing “evolutionary” perturbations in the key components thatcontribute to the biomimetic behavior and measuring the transfer energythat arises from those evolutionary perturbations. The transfer energycan be estimated using biomechanical models to augment the inertialmeasurement unit feedback, or, in special cases, temporary inertialmeasurement unit subsystems (an IMU mounted on the body in the form of abelt around the torso and/or upper leg) could be used to facilitateestimate of torso pose and body center-of-mass velocity. Using theFletcher-Powell method (or other suitable optimization method known tothose skilled in the art), an intelligent evolution of parameters can beintroduced and an optimum can be calculated. This optimum could, due tothe rehabilitative effects of the augmentation, change over time. Byapplying these evolutionary perturbations continually and slowly overtime, the optimum can be achieved on a continual basis. Or, as would bethe case at the initial fitting or medical checkup of the prosthesis ororthosis, this evolutionary optimization could occur over a much shorterinterval, say, in five to ten minutes.

The following is a description of the different phases of a subject'sgait cycle and, in one embodiment, the steps performed by a ankle jointprosthesis according to principles of the invention are for sensing theoperation of and for controlling the ankle joint prosthesis.

Controlled Plantarflexion

At impact, check that the ground reaction force and the zero momentpivot correspond with the part of the foot that we expect (from theterrain discrimination model) to hit the ground first. Confirm thatthere is a corresponding change in the ankle angle (or ankle torque) andthat the appropriate end of the foot is stationary. After impact lookfor a condition where the local terrain slope corresponding to theinertial foot-flat angle is significantly less than expected. Saturateankle spring restoring force and increase damping when this is detected.For terrain discrimination, based upon the biomechanical model feedbackconfirm that the terrain hypothesis (slope vs. stair) is correct andthat the wearer hasn't tripped. For example, a tripping event on a stairmight be detected as a large negative force in the y-direction insteadof a large z-force centered on the forward part of the foot. For terraintexture, either the heel or the forward part of the foot will impactfirst. The non-elastic component of the depression associated with thisimpact will be computed. On hard ground, this depression should benegligible—only an elastic deformation (foot module, linear actuator)will be observed. In mud or soft ground, terrain plasticity will beobserved by looking at the trajectory of the impacting foot segment. Theterrain plasticity will be used as an attenuator on the net work that isperformed on this walking cycle. Slipping can also be detected by notingthe forward velocity of the impacting foot segment after impact. Anescalator or people mover can be detected by noting that the shank angleis not rotating in accordance with the forward velocity of the foot,signaling that the wearer is well balanced and is stepping onto a movingsurface. For impedance control of the ankle joint apply optimalimpedance using estimated terrain-referenced velocity attack angle (y)lower-limb momentum, estimated terrain slope and terrain property. Forreflex control, in the event that slipping is detected, abalance-restoring reflex will be generated to move the knee over theankle. For balance control, optimal balance will normally be achieved byinertially referencing the spring equilibrium after the local terrainslope estimate is updated at foot-flat. In the event that the terrain isslippery, the algorithms that maintain balance will introduce a positivetorque “reflex” to “pull” the shank forward in order to assist thewearer as she works to position the knee over the ankle—thereby gettingthe body center-of-mass aligned with the estimated ground reactionforce.

Controlled Dorsiflexion

Once foot flat is detected, the controller inertially references thespring equilibrium angle for this local terrain slope so that when thewearer is standing in alignment with gravity on this slope, no restoringtorque is applied by the ankle under static conditions. At this point,the local terrain context is now known precisely. Foot referencecoordinates at this “foot flat” position are also defined for use inassessing the impact of terrain texture. For terrain texture, thealgorithms use integrated measures of slip and deformation relative tothe “foot flat” reference to update the terrain propertymodel—specifically to measure plasticity of the surface and it'sslipperiness by measuring how the impacted foot segment moves betweenfoot-strike and foot flat. These measures can be used to attenuate ankleimpedance and net work (reflex torque in late plantar flexion. Also, if“slipping” is detected between foot-strike and foot-flat, an algorithmimplemented in the controller, also looks at shank angular velocity (howthe knee is moving in relation to the ankle joint) to discriminatebetween a slippery surface and an escalator/people-mover. In eitherevent, the zero-velocity update would not be scheduled since no reliable“ankle joint at zero velocity” will be available on this step. In theevent that the terrain is slippery, special measures will need to beinvoked by the balance function. In the case where the foot lands on amoving escalator or people mover, nominal impedance can be used on thenew inertial frame. For impedance control, the control system can applyoptimal impedance that maintains an inertially-referenced equilibriumangle; creates a walking speed-dependent stiffness (lower stiffness forfaster walking speed) to enable a higher level of net work; and reducesthe stiffness in slippery or highly-plastic surfaces. For reflexcontrol, in the event that slipping is detected, a balance-restoringreflex will be generated to move the knee over the ankle. For balancecontrol, optimal balance will normally be achieved by inertiallyreferencing the spring equilibrium after the local terrain slopeestimate is updated at foot-flat. In the event that the terrain isslippery, the algorithms that maintain balance will introduce a positivetorque “reflex” to “pull” the shank forward in order to assist thewearer as she works to position the knee over the ankle—thereby gettingthe body center-of-mass aligned with the estimated ground reactionforce.

Powered Plantarflexion

The model monitors slippage and sinking into the surface and identifiesankle torque limits that can be used to make ambulation efficient inthese conditions. For terrain texture, terrain property estimates arerefined in this state and are used as an input to the impedance, reflexand balance functions. For impedance control, nominal impedanceparameters will be modified to accommodate changes in walking speed,terrain surface characteristics and deformation and foot slippage. Aspecial “force field”—typically a non-linear actuator force thatexponentially increases as the ball-nut approaches a predefined end-stoplimit—is applied by the motor controller to make sure that the K3 springenergy (in the parallel elastic member) does not exceed the lower boundof its fracture limit. For reflex control, reflex amplitude will beadjusted to account for the net work “setpoint” from the biomechanicalmodels in combination with the degree to which the terrain can supportproduction of this net work. For balance control, optimal balance willnormally be achieved by inertially referencing the spring equilibriumafter the local terrain slope estimate is updated at foot-flat. In theevent that the terrain is slippery, the algorithms that maintain balancewill introduce a positive torque “reflex” to “pull” the shank forward inorder to assist the wearer as she works to position the knee over theankle—thereby getting the body center-of-mass aligned with the estimatedground reaction force.

Early Swing

For early swing, shortly after the toe leaves the ground, the modelmonitors the inertial trajectory of the ankle, heel and toe anddetermines when the ankle can be dorsiflexed back to its neutralposition without being obstructed by the terrain. The model computes anoptimal trajectory with suitable impedance gains and feed-forward torqueto move the ankle to the neutral position (to avoid tripping hazards) inthe quickest, efficient and stable fashion. For terrain discrimination,the model starts to keep track of the swept (“no contact” with footmember) volume through which the foot has moved thereby informing theadaptive ankle positioning function in late swing when a toe-downsolution is the only viable solution (e.g., to land on a shallow stairor ledge). For impedance control in early swing, a neutral value ofimpedance is applied by the controller. A force-field function isapplied to make sure that the linear actuator does not impact the hardstop (end of travel)—a condition that could cause the actuator to stickthere (at the end of travel). For impedance control in early swinginformed by the hybrid biomechanical model, the controller controls theimpedance to create a trajectory that exponentially drives theequilibrium position (ankle angle setpoint) to the desired neutralposition. A feed-forward torque function is applied to reduce theinteraction between impedance characteristics and the ankle anglefollowing error that could otherwise introduce overshoot and ringing,for instance.

Late Swing

For terrain discrimination, the model keeps track of the “clear” volumethrough which the foot has moved thereby informing the adaptive anklepositioning function in late swing when a toe-down solution is the onlyviable solution, say, to land on a shallow stair or ledge. Moregenerally, the ankle trajectory is monitored and pattern recognitionfunctions are used to determine the likelihood that the foot will belanding on a stair/ledge as opposed to a sloping surface. One simple waythat we have found to discriminate between the two conditions is tomeasure the angle that the ankle velocity makes in relation to vertical;where in various experiments it was determined that when this angle isless than 10 degrees, the foot will land on a horizontal step. Forimpedance control, informed by the terrain discrimination model, theankle trajectory (equilibrium) will be modified by the controller asneeded to avoid tripping hazards. For example, if the terraindiscrimination function assigns the maximum likelihood to stair ascent,additional dorsiflexion may be commanded to make sure that the toe doesnot catch on the stair or ledge. As before, the hybrid biomechanicalmodel plans a continuously updatable equilibrium trajectory that can befollowed safely and in a stable fashion with tight tolerances. In thelate-stance state, the biomechanical model computes the optimumequilibrium angle and ankle impedance that will minimize an objectivefunction that includes some combination of transfer energy and knee-hipimpact forces. This optimization function could be implemented via tablelookup in the State Machine ROM. Or, in the preferred embodiment, theState Controller function will perform the optimization in real-time,using approximations of the rigid-body dynamics, to compute and optimizethe objective functions.

Variations, modifications, and other implementations of what isdescribed herein will occur to those of ordinary skill in the artwithout departing from the spirit and the scope of the invention asclaimed. Accordingly, the invention is to be defined not by thepreceding illustrative description but instead by the spirit and scopeof the following claims.

The invention claimed is:
 1. A motorized ankle prosthesis, orthosis orexoskeleton apparatus, comprising: an actuator operatively coupled to ajoint; at least one sensor for determining a parameter that correlateswith gait speed; a controller in communication with the at least onesensor, the controller configured to cause the actuator to deliver tothe joint over a gait cycle an amount of net ankle work that increaseswith increasing gait speed, the controller further configured tocalculate net ankle work applied to the joint over the gait cycle; andan interface in communication with the controller, the interfaceconfigured to display information regarding the calculated net anklework applied to the joint, to receive an input from an external sourceand send a signal based on the received input that causes the controllerto change net ankle work over the gait cycle applied to the joint as afunction of gait speed.
 2. The apparatus of claim 1, wherein thecontroller is configured to apply torque to the joint based on a signalfrom the interface, the applied torque comprising at least one of ajoint impedance, a joint equilibrium, a joint damping and an anklereflex response.
 3. The apparatus of claim 2, wherein the ankle reflexresponse comprises at least one of a positive torque feedback responseand a positive velocity feedback response.
 4. The apparatus of claim 3,wherein the torque applied is represented by the followingrelationships:Γ=−(k(θ−θ₀)+b{dot over (θ)}+j{umlaut over (θ)})+Γ_(reflex)(α)Γ_(re flex)(α)=Γ_(Torque reflex)(α)+Γ_(Velocity reflex)(α) whereΓ=torque applied, k=joint impedance, θ₀=joint equilibrium, b=jointdamping, j=joint inertia, Γ_(reflex)(α)=ankle reflex response,Γ_(Torque re flex)(α)=positive torque feedback response,Γ_(Velocity reflex)(α)=positive velocity feedback response, and α=one ormore parameters that defines a torque response.
 5. The apparatus ofclaim 4, wherein a comprises an exponential parameter such that thepositive torque feedback response relates exponentially to a measuredtorque, and the positive velocity feedback response relatesexponentially to an angular velocity.
 6. The apparatus of claim 1,wherein the speed of a point located on or coupled to the apparatus isat least one of a translational speed and an angular speed.
 7. Theapparatus of claim 1, wherein the joint is a mechanical joint or abiological joint.
 8. The apparatus of claim 1, wherein the controller isconfigured to adjust ankle stiffness based on an indication from theinterface, during at least part of a stance or swing phase.
 9. Theapparatus of claim 1, wherein the interface is in wireless communicationwith the controller.
 10. The apparatus of claim 1, wherein the interfaceis configured to tune the controller such that an ankle joint responseconforms to a biomimetic norm determined by a biomechanical datacollection.
 11. The apparatus of claim 10, wherein the biomimetic normcomprises at least one of a biological level of net work, stiffness,damping, and power.
 12. The apparatus of claim 11, wherein thebiomimetic norm is a function of walking speed.
 13. The apparatus ofclaim 1, wherein the interface is configured to tune the controller tocause the actuator to adjust at least one of a Controlled PlantarFlexion stiffness, a Controlled Dorsiflexion stiffness and powerdelivered during Powered Plantar Flexion.
 14. The apparatus of claim 1,wherein the at least one sensor comprises at least one of an angularrate sensor, an accelerometer, a displacement sensor and a sensor fordetermining an inertial pose of a portion of the apparatus.
 15. Theapparatus of claim 1, wherein the controller is configured to cause theactuator to decrease stiffness of the joint with increasing walkingspeed during the gait cycle.
 16. The apparatus of claim 1, furthercomprising a spring operatively coupled to the actuator and the joint.17. The apparatus of claim 1, wherein the controller is configured tocause the actuator to deliver net positive ankle work to the joint basedon a signal from the interface during the gait cycle.
 18. The apparatusof claim 1, further comprising an inertial measurement unit, fordetermining at least one of ambulation speed of the wearer andunderlying terrain.
 19. A method of operating a motorized ankleprosthesis, orthosis or exoskeleton apparatus, comprising: inputtinginformation regarding wearer need to an interface, wherein the interfaceis in communication with a controller, the controller configured tocause an actuator, operatively coupled to a joint, to deliver an amountof net ankle work to the joint over a gait cycle, the controller furtherconfigured to calculate net ankle work applied to the joint over thegait cycle; displaying on the interface information related to thecalculated net ankle work applied to the joint; and adjusting the netankle work over the gait cycle applied to the joint by tuning thecontroller based on the information input to the interface.
 20. Themethod of claim 19, wherein net ankle work is adjusted across gaitspeed.
 21. The method of claim 19, wherein tuning the controllercomprises wireless communication between the interface and thecontroller.
 22. The method of claim 19, wherein the information input tothe interface tunes the controller such that an ankle response conformsto a biomimetic norm determined by a biomechanical data collection. 23.The method of claim 22, wherein the biomimetic norm comprises at leastone of a biological level of net positive work, stiffness, damping andpower.
 24. The method of claim 23, wherein the biomimetic norm isadjusted as a function of walking speed.
 25. The method of claim 19,wherein tuning the controller comprises adjusting a response of theactuator to achieve at least one of a Controlled Plantar Flexionstiffness, a Controlled Dorsiflexion stiffness and power deliveredduring Powered Plantar Flexion.
 26. The method of claim 19, furthercomprising sensing at least one of a translational and angular speed ofa location on or coupled to the apparatus during the gait cycle.
 27. Themethod of claim 26, wherein the location coupled to the apparatuscomprises a location on a wearer coupled to the apparatus.
 28. Themethod of claim 27, further comprising determining an inertial posetrajectory comprising at least one of a) an orientation component of atleast one location on or coupled to the apparatus, b) a translationcomponent of the at least one location on or coupled to the apparatus,and c) a velocity component of the at least one location on or coupledto the apparatus.